X-ray collimation assembly

ABSTRACT

An x-ray imaging system according to the present invention comprising a stepped scanning-beam x-ray source and a multi-detector array. The output of the multi-detector array is input to an image reconstruction engine which combines the outputs of the multiple detectors over selected steps of the x-ray beam to generate an x-ray image of the object. A collimating element, preferably in the form of a perforated grid containing an array of apertures, interposed between the x-ray source and an object to be x-rayed. A maneuverable positioner incorporating an x-ray sensitive marker allowing the determination of the precise position coordinates of the maneuverable positioner.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. Pat. Ser. No. 08/419,740filed on Apr. 10, 1995, which issued as U.S. Pat. No. 5,610,967, whichis a continuation of U.S. patent application Ser. No. 08/386,861, filedFeb. 10, 1995, which issued as U.S. Pat. No. 5,651,047, which acontinuation-in-part of U.S. patent application Ser. No. 375,501, filedJan. 17, 1995 now abandoned, which is a continuation of U.S. patentapplication Ser. No. 08/042,742, filed Apr. 5, 1993, now abandoned; ofU.S. patent application Ser. No. 08/342,641, filed Nov. 21, 1994, nowabandoned which is a continuation of U.S. patent application Ser. No.08/008,455, filed Jan. 25, 1993, now abandoned; and, of Internationalpatent application Ser. No. PCT/US94/03737, filed Apr. 5, 1994, whichdesignated the United States from which priority is claimed under theprovisions of 35 U.S.C. §§120 and 365, all of which are incorporatedherein by reference in their entirety, each of which are incorporatedherein by reference in their entirety.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The field of the present invention pertains to diagnostic x-ray imagingequipment. More particularly, the present invention pertains toreal-time scanning-beam x-ray imaging systems.

2. Description of Related Art

Real-time x-ray imaging is increasingly being required by medicalprocedures as therapeutic technologies advance. For example, manyelectro-physiologic cardiac procedures, peripheral vascular procedures,PTCA procedures (percutaneous transluminal catheter angioplasty),urological procedures, and orthopedic procedures rely on real-time x-rayimaging. In addition, modern medical procedures often require the use ofinstruments, such as catheters, that are inserted into the human body.These medical procedures often require the ability to discern the exactlocation of instruments that are inserted within the human body, oftenin conjunction with an accurate image of the surrounding body throughthe use of x-ray imaging.

Current clinical real-time x-ray equipment produces high levels of x-rayexposure to both patients and attending staff. The United States Foodand Drug Administration (F.D.A.) has reported anecdotal evidence ofacute radiation sickness in patients, and concern among physicians ofexcessive occupational exposure. (Radiological Health Bulletin, Vol.XXV1, No. 8, August 1992).

A number of real-time x-ray imaging systems are known. These includefluoroscope-based systems where x-rays are projected into an object tobe x-rayed and shadows caused by relatively x-ray opaque matter withinthe object are displayed on the fluoroscope located on the opposite sideof the object from the x-ray source. Scanning x-ray tubes have beenknown in conjunction with the fluoroscopy art since at least the early1950s. Moon, Amplifying and Intensifying the Fluoroscopic Image by Meansof a Scanning X-ray Tube, Science, Oct. 6, 1950, pp. 389-395.

Reverse-geometry scanning-beam x-ray imaging systems are also known. Insuch systems, an x-ray tube is employed to generate x-ray radiation.Within the x-ray tube, an electron beam is generated and focussed upon asmall spot on the relatively large anode (transmission target) of thetube, inducing x-ray radiation emission from that spot. The electronbeam is deflected (electromagnetically or electrostatically) in a rasterscan pattern over the anode target. A small x-ray detector is placed ata distance from the anode target of the x-ray tube. The detectortypically converts x-rays which strike it into an electrical signal inproportion to the detected x-ray flux. When an object is placed betweenthe x-ray tube and the detector, x-rays are attenuated and scattered bythe object in proportion to the x-ray density of the object. While thex-ray tube is in the scanning mode, the signal from the detector isinversely proportional to the x-ray density of the object.

Examples of known reverse-geometry scanning-beam x-ray systems includethose described in U.S. Pat. No. 3,949,229 to Albert; U.S. Pat. No.4,032,787 to Albert; U.S. Pat. No. 4,057,745 to Albert; U.S. Pat. No.4,144,457 to Albert; U.S. Pat. No. 4,149,076 to Albert; U.S. Pat. No.4,196,351 to Albert; U.S. Pat. No. 4,259,582 to Albert; U.S. Pat. No.4,259,583 to Albert; U.S. Pat. No. 4,288,697 to Albert; U.S. Pat. No.4,321,473 to Albert; U.S. Pat. No. 4,323,779 to Albert; U.S. Pat. No.4,465,540 to Albert; U.S. Pat. No. 4,519,092 to Albert; and U.S. Pat.No. 4,730,350 to Albert.

In a typical known embodiment of a reverse-geometry scanning-beamsystem, an output signal from the detector is applied to the z-axis(luminance) input of a video monitor. This signal modulates thebrightness of the viewing screen. The x and y inputs to the videomonitor are typically derived from the signal that effects deflection ofthe electron beam of the x-ray tube. Therefore, the luminance of a pointon the viewing screen is inversely proportional to the absorption ofx-rays passing from the source, through the object, to the detector.

Medical x-ray systems are usually operated at the lowest possible x-rayexposure level at the entrance of the patient that is consistent withthe image quality requirements (particularly contrast resolution andspatial resolution requirements) for the procedure and the system.Typical patient entrance exposure in conventional 9" field of view imageintensifier systems used in cardiac procedures, in the AP (anteriorposterior) view with a standard adult chest, is approximately 2.0 to 2.8R/min. The term "low dosage" used herein refers to a factor of 2 to 20less than this.

Time and area distributions of x-ray flux follow a Poisson distributionand have an associated randomness which is unavoidable. The randomnessis typically expressed as the standard deviation of the mean flux, andequals its square root. The signal-to-noise ratio of an x-ray imageunder these conditions is equal to the mean flux divided by the squareroot of the mean flux. i.e., for a mean flux of 100 photons, the noiseis ±10 photons, and the signal-to-noise ratio is 10.

Accordingly, the spatial resolution and the signal-to-noise ratio ofx-ray images formed by known reverse-geometry scanning x-ray imagingsystems are dependent, to a large extent, upon the size of the sensitivearea of the detector. If the detector aperture is increased in area,more of the diverging rays are detected, effectively increasingsensitivity and improving the signal-to-noise ratio. At the same time,however, the larger detector aperture reduces attainable spatialresolution as the "pixel" size (measured at the plane of the object tobe imaged) becomes larger. This is necessarily so because most objectsto be imaged in medical applications (e.g., structures internal to thehuman body) are some distance from the x-ray source. In the knownsystems, therefore, the detector aperture size has been selected so asto effect a compromise between resolution and sensitivity, it not beingpreviously possible to maximize both resolution and sensitivitysimultaneously.

In the medical field, several conflicting factors, among them patientdosage, frame rate (the number of times per second that the object isscanned and the image refreshed), and resolution of the image of theobject, often work to limit the usefulness of an x-ray imaging system.For example, a high x-ray flux may easily yield high resolution and ahigh frame rate, yet result in an unacceptably high x-ray dosage to thepatient and attending staff.

Similarly, lower dosages may be achieved from the known systems at thecost of a low resolution image or an inadequate refresh rate. Apreferred medical imaging system should provide low patient dosage, highresolution and an adequate refresh rate of up to at least about 15images per second--all at the same time. Therefore, systems such as theknown reverse-geometry scanning-beam x-ray imaging systems describedabove are not acceptable for diagnostic medical procedures whereexposure times are relatively long and where, as is always the case withlive patients, the x-ray dose received by the patient should be kept toa minimum.

Minimally invasive procedures in medicine are typically characterized byaccess to areas inside the body using existing orifices such as theureter or by percutaneous entry such as a puncture of the femoral vein.In such procedures, various tools and catheters may then be progressedinto the body and maneuvered using a real-time x-ray imaging system forguidance. An estimated 3,000,000 medical procedures of this type wereperformed in 1993 under x-ray fluoroscopy guidance. Many of theseprocedures involve the introduction of a catheter into the coronaryarteries and the heart, and the evaluation of cardiac function byinspection of images taken when contrast media is introduced via a lumenin the catheter. Some of the tools that may be inserted in this mannerinclude lasers where the laser device is located outside the body andthe laser light delivered to the site of interest with a fiber-opticwave guide disposed in a catheter, drug delivery systems adapted todeliver precisely measured quantities of a specific drug or radiologicalmaterial to the site of interest, ultrasound systems in which atransducer on the tip is used to view a site of interest by deliveringthe image over to a video system which can then display and recordimages of the site of interest, and other tools known to the art. It isalso possible to adapt such procedures to non-medical applications whereaccess is difficult and the value of the procedure high, e.g., enginediagnosis and repair.

As used herein, the term "maneuverable positioner" is meant tocollectively include and refer to, for example, catheters, probes,endoscopes, and other maneuverable positioners and tools.

The known medical x-ray imaging devices do not provide a highly-accuratedetermination of location for maneuverable positioners with a preciseimage of the patient's internal structure. Generally, the physicianusing known systems can roughly ascertain the position of maneuverablepositioners relative to body features within the patient, but precisionand repeatability, the ability to return to the exact same place,especially in the axis parallel to the x-ray beam, is lacking. Thus thedistance between the x-ray emitting source and the maneuverablepositioner within the body may not be readily or accurately determinedwith the precision useful in today's advanced medical procedures, whichmay require, among other things, the ability to determine a positionwith the maneuverable positioner, move the maneuverable positioner, andreturn the maneuverable positioner to the exact same place.

For example, since 1982 there has been increasing use of catheterablation to cure certain types of arrhythmia. In these types ofarrhythmia, such as Wolff-Parkinson-White syndrome, the conductivecongenital muscle fibers can be made nonconductive by heating themlocally to a sufficient temperature to cause scar tissue to form. Mostof these ablations are done with radio-frequency energy but the emittingelectrode must be placed within one to three millimeters of the musclefiber location and it must stay in intimate contact with it for a numberof heartbeats and respiratory cycles.

Although the treatment of arrhythmia through catheter ablation has someadvantages, there are also some problems. The advantages of theprocedure are that it has a very high success rate, it is minimallyinvasive, it can be performed in a few hours in a procedure room, and itis considerably less expensive than open chest surgery or a lifetime ofdrug therapy. The major disadvantage is that the length of the procedureis uncertain and typically long. This leads to difficulty in schedulingphysicians and facilities, fatigue for both patient and staff, andhigh-radiation dosages for patient and physician.

Attempts to solve these problems have focused mainly on providing moresteerable catheters to reduce the time to find the precise location ofthe ablation site and to position the catheter for remaining in contactwith the substrate during the ablation time, which is typically five toninety seconds. Having more steerable catheters has not yet reduced thetime or uncertainty of time because the location of the catheter isgenerally determined by looking at an x-ray image projected on a monitorand by analyzing the electrocardiogram. Both of these actions must bedone in real time in order to know whether to move the catheter and inwhich direction to move it. The actual direction of movement may beuncertain due to the nature of an x-ray image of soft tissue and blood,the poor control and feedback of the catheter, the movement of theheart, and the difficulty of determining direction from theelectrocardiogram analysis.

In the U.S., there are currently 300,000 to 500,000 people who die eachyear due to arrhythmia that is a result of a myocardial infarction.However, it is believed that if the slow-conduction zone around theinfarct could be electrically mapped and selectively ablated, that acure could be obtained. Tests on animals and some humans havedemonstrated the possibility of such a procedure but the success ratehas been low. The reason for the low success is thought to be the needto map the entire area of the infarct and slow conduction zone and thento be able to ablate multiple sites without depending on acquiring acharacteristic electrogram once the ablation has begun. Currentinvestigations attempting to solve the problem utilizing a catheternetwork array of nodes suffers from the problem of extracting thecatheter network array from inside the heart without damaging theinternal structure of the heart.

For various reasons, the imaging modalities of MRI, CT, and ultrasoundare not normally suitable when anatomical markers are needed duringcardiac diagnostic and treatment procedures. In addition, the use ofknown methods employing x-ray fluoroscopes for imaging typically has theserious disadvantage of not being able to distinguish anatomical detailinside the heart. The physician relies on the shadows generated, his orher intimate knowledge of the anatomy, the characteristic movement ofthe image and catheters caused by the cardiac cycle and the respiratorycycle, and for fine positioning, the electrocardiogram.

Accordingly, there is a need for devices and methods to provide aprecise determination of the coordinates of a maneuverable positionerwithin a human patient during a medical procedure. The same techniquesand apparatus can also be used to advantage in any x-ray procedure whichrequires accurate determination of the X, Y and Z coordinates of theposition of a maneuverable positioner which may be adapted to sensex-rays.

SUMMARY OF THE INVENTION

An x-ray imaging system according to the present invention comprises ascanning-beam x-ray source and a multi-detector array. The output of themulti-detector array is input to an image reconstruction engine whichcombines the outputs of the multiple detectors over selected positionsof the x-ray beam to generate a real-time x-ray image of the object.

An embodiment of an aspect of the invention includes an x-ray tubeincluding a charged particle beam source and an anode target. Beamcontrol circuitry focusses the charged particle beam and directs orscans the beam across the anode target in a predetermined pattern. Forexample, the predetermined pattern may be a raster scan pattern, aserpentine or "S" shaped pattern, a spiral pattern, a random pattern, agaussian distribution pattern centered on a predetermined point of theanode, or such other pattern as may be useful to the task at hand.

A collimating element, preferably in the form of a grid, may beinterposed between the x-ray tube and an object to be x-rayed. In onepreferred embodiment, the collimating element is composed of a roundmetal plate having a diameter of about 25.4 cm (10 in) and includes astaggered array of apertures numbering 500 by 500 at the center row andcolumn of the collimating element. The collimating element is preferablyplaced immediately in front of the emitting face of the x-ray tube.Other collimating element configurations may also be used. In onepreferred embodiment, each of the apertures in the collimating elementis constructed so that each of the axes of each of the apertures isdirected toward (or points at) a detection point, e.g., the center of amulti-detector array, located a selected distance from the collimatingelement. That distance is selected to allow placement of the object tobe x-rayed between the collimating element and the multi-detector array.In the preferred embodiment, the function of the collimating element isto form thin pencil beams of x-rays, all directed from a focal spot onthe anode target of the x-ray tube toward the multi-detector array.

A multi-detector array, preferably containing an array of detectorelements (preferably an area array such as a DET_(x) by DET_(y)rectangle or square, or, more preferably, a pseudo-round array), iscentered at the detection point. The multi-detector array preferablycomprises a plurality of densely packed x-ray detectors. Themulti-detector array is designed, positioned and applied, according tothe present invention, in a manner that yields high sensitivity withoutloss of resolution. This results in an x-ray system having a resolutioncomparable to or better than that of known conventional x-ray systems atan exposure at least an order of magnitude less than that of the knownx-ray systems. This aspect of the present invention provides importantbenefits in medical and other applications. X-ray dosage to patients andattending medical staff is reduced when using this aspect to performcurrent medical procedures. Procedures now believed to have too high aradiation exposure risk may become acceptable.

The output of the multi-detector array is preferably an intensity valuefor each detector of the multi-detector array for each x-ray beamemitted through an aperture in the collimating element. Because eachaperture is located at a different point in space relative to themulti-detector array and the object under investigation, differentoutputs will be available from each detector of the multi-detector arrayfor each aperture that the x-ray beam travels through. Themulti-detector array output may be converted into an image in a numberof ways.

The imaging system of the present invention is also capable of use instereo imaging. In one embodiment, the collimation element contains twogroups of apertures. For stereo imaging, the axes of one group ofapertures is constructed to point to a first detection point on a firstmulti-detector array and the axes of a second group of apertures isconstructed to point to a second detection point on a secondmulti-detector array. By constructing two images from the outputs of themulti-detector array and using conventional stereoscopic displaymethods, a stereo image may be produced.

An imaging system of the present invention is also capable ofhighlighted imaging of materials which exhibit different x-raytransmissivities at different x-ray photon energies. Accordingly, forexample, microcalcification, which is associated with approximately 60%of the breast cancer diagnosed, may be imaged. Calcium is also typicallyassociated with heart disease when found in the coronary arteries. Inone embodiment, by constructing the collimation element and/or anodetarget to sequentially emit two or more groups of x-rays beams eachhaving different x-ray energy spectra, and directing each group to themulti-detector array (more than one multi-detector arrays could also beused), the difference of transmissivities of the object underinvestigation at the various x-ray photon energies can be used to createan image, thus highlighting only those materials within the object underinvestigation which exhibit differential x-ray transmissivity. Optimizedfor the detection of calcium, for example, such an imaging system is apowerful tool for use in the early detection of breast cancer and otheranomalies.

Utilizing a multi-detector array which intercepts the entire x-ray beamemitted from each aperture of the collimator element and imageprocessing the array output is the preferred embodiment of the detector.It provides a maximum sensitivity without sacrificing the resolutionprovided by using a single small area detector. While a single detectorof the same area as the multi-detector array would provide the samesensitivity, it would do so at the cost of a loss of resolution.

Additionally, sampling techniques utilizing information from less than a1:1 image pixel to aperture ratio may be used for generating data fromthe multi-detector array which can reduce the complexity of the system,required processing speed, and energy consumption while providingvirtually the same image quality.

An aspect of the present invention can also be used to identify theunique location of a marker transported within another object by amaneuverable positioner. In its most general sense, this would beaccomplished by an x-ray sensitive marker disposed in a body andincludes the transmission of an indication of the presence of x-rayradiation outside of the body in which it is disposed.

According to one embodiment of this aspect of the invention useful inmedical applications, a catheter comprises an elongated body having adistal end adapted to be inserted into a body cavity, blood vessel,digestive tract, or the like and a proximal end available to a personperforming a medical procedure. The catheter includes at least one lumenrunning therethrough. An optical fiber is disposed in the lumen andextends from the distal end to the proximal thereof. A miniaturized("mini") x-ray sensor comprised of an x-ray sensitive material isdisposed at the end of the optical fiber positioned at the distal end ofthe catheter. The end of the optical fiber at the proximal end of thecatheter is coupled to a photodetector. The reaction of the sensormaterial of the mini x-ray sensor to an x-ray beam sequentiallytransmitted through a collimator, coupled with the transmission of thatreaction to the photodetector, allows the determination of the preciseposition of the sensor material. Embodiments of the techniques todetermine this precise position from the mini x-ray sensor reaction isdiscussed in detail below.

Another aspect of the present invention is the ability to determine thedistance of a maneuverable positioner containing a mini x-ray sensorfrom a known reference plane. Each x-ray beam emitted through acollimation grid aperture of the present invention is shaped like adiverging cone with its apex at the anode target and its divergenceangle determined by electron beam spot size on the anode target and thegeometry of the collimation grid apertures with respect to the anodetarget. The divergent beams are designed to overlap more and more thefarther you get from the x-ray source. The mini x-ray sensor in thisaspect is preferably disposed in a maneuverable positioner and may have(but is not required to have) a size smaller than the spacing betweenthe apertures of the collimation grid. When such a sensor is disposed inthe x-ray field it will detect, during a complete scan cycle, x-raysfrom only a certain number of apertures, the number depending upon themini x-ray sensor's distance from the output face of the collimationgrid. When the mini x-ray sensor is located close to the output face ofthe collimation grid, it will react to x-ray pulses from a first numberof apertures per scan cycle. At a greater distance from the output face,it will react to x-ray pulses from a second number of apertures greaterthan the first number. When the mini x-ray sensor is near the x-raymulti-detector array, it will react to x-ray pulses from an even greaternumber of apertures per scan cycle. By calibrating the number ofapertures per scan cycle to which the mini x-ray sensors reacts with themini x-ray sensor's distance from a known reference, the distance of themini x-ray sensor from the reference may be determined by consulting alook-up table and/or by interpolation.

When used as described herein, the above-described embodiment of thepresent invention answers the long felt need for anatomical markersduring cardiac diagnostic and treatment procedures. Since typicalreference positions for electrocardiograms are the high right atrium,the bundle of HIS, the apex of the right ventricle, and the coronarysinus, there is an opportunity to have at least three points located inthe x-ray image during the cardiac cycle. These three points canprecisely locate the coordinates of the ablation catheter. Knowing thecoordinates of these points, the physician can then map an area with thecatheter and correlate its position with the electrocardiograms. He canthen return to the same spot after leaving it and can measure bounce orother movement, he can also measure internal dimensions of the heartmapping points, determine wall thicknesses, build 3D images from thedata and overlay cardiac action Potentials. In addition, in a subsequentprocedure the same locations may be found from overlaying the maps ofanatomy and cardiac potentials.

The initial work performed with this aspect of the invention was toanalyze the images from a biplane x-ray system when it was gated to thecardiac cycle. In some cases this positioning using image analysis isadequate but the precision can be greatly improved if an x-ray marker isincluded in the catheter. With conventional image intensifiertechnology, such a point sensor would not be useful since the entirefield of view is irradiated simultaneously. However, in a scanning-beamx-ray system, the beam irradiates only a small field of view at a giventime and therefore the location of the sensor in each individualcatheter can be uniquely identified. By utilizing a stereo or biplanescanning-beam system, the sensor can be located in three dimensionsprovided that the two beams are synchronized. The advantage of the abovedescribed medical catheter embodiment of the invention is that itpermits existing catheters inside the patient to now also function asanatomical markers when a mini x-ray sensor is employed, significantlyreducing the time to map and ablate. Additional advantages are moredetailed mapping of the cardiac substrate, correlation of theintercardiac electrodes with anatomical location, display in threedimensions of the intercardiac electrodes on an image of the heart, andcomparison of electrograms from studies done at different times byoverlaying.

It is an object of one aspect of the present invention to provide ascanning-beam x-ray imaging system capable of use in medical diagnosticprocedures undertaken on living human patients.

It is also an object of another aspect of the present invention toprovide a scanning-beam x-ray imaging system which provides highresolution images at adequate frame rates while minimally exposing theobject under investigation to x-ray radiation.

It is a further object of another aspect of the present invention toprovide a scanning-beam x-ray imaging system having improved resolutionat a distance from the plane of the source of the x-rays whilemaintaining decreased x-ray flux levels.

It is also a further object of an aspect of the present invention toprovide a method and apparatus for precisely determining the position ofa maneuverable positioner within an object undergoing an x-rayprocedure.

It is yet a further object of an aspect of the present invention toprovide a method and apparatus for precisely and simultaneouslydetermining and displaying information related to the X, Y and Zcoordinates of a maneuverable positioner within an object undergoing anx-ray procedure.

It is also a further object of an aspect of the present invention toprovide an electronic glove and other improvements in safety for medicalapplications by feedback of position from a mini x-ray sensor.

It is an object of another aspect of the present invention to providefor improved image quality by employing region of interest scanning.

It is a further object of an aspect of the present invention to providea scanning beam x-ray imaging system for non-medical applications wherescatter may degrade image quality, e.g., to image or inspect honeycombairplane structures, corrosion, and printed circuit boards.

An advantage of an aspect of the present invention is that it canprovide a method and apparatus for generating a "road map." For example,once a maneuverable positioner incorporating a mini x-ray sensor inaccordance with the present invention has been threaded to a particularlocation of interest within the body or object, it can be removed andthen re-threaded along the same path by generating waypoints, e.g., bydetermining X, Y and Z coordinates of various locations passed throughduring the first insertion. These waypoints may be obtained asfrequently as desired along the first path taken to the location ofinterest to facilitate retracing the same path on subsequent occasions.This aspect of the present invention may have important application inintravascular and intracardiac ultra sound procedures.

Another advantage of an aspect of the present invention is that it canprovide a method and apparatus to precisely locate and monitor the shapeor position of stents. The precise location of the surface of a stentcan be obtained through use of an x-ray sensing maneuverable positionerin accordance with the present invention by defining points along thesurface of the stent, determining the X, Y and Z coordinates of thosepoints and recording them. Subsequently, X, Y, and Z coordinates forthose defined points can be redetermined and recorded over time andchanges in shape or position of the stent can be observed and plotted.

Another advantage of an aspect of the present invention is that it canprovide a method and apparatus for repeatable delivery of drugs,radiologic and similar materials to a specific site in the body.

These and many other objects and advantages of the present inventionwill become apparent to those of ordinary skill in the art from aconsideration of the drawings and the description of the inventioncontained herein. The principles of the present invention may beemployed in any application, medical or industrial. Principles oraspects of the present invention can be applied for example wherelocation of internal features of an object is desired and insertion ofan x-ray sensitive device is feasible. Industrial applications arevariously called x-ray inspection, x-ray analysis, failure analysis,non-destructive testing, and in-situ testing.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a diagram showing the basic components of a preferred lowdosage scanning-beam x-ray imaging system.

FIG. 2 is a diagram showing the distribution of x-rays in the forwarddirection from a scanning-beam x-ray imaging system in the absence of acollimation grid.

FIG. 3 is an enlarged cross sectional representation of a portion of apreferred collimation grid and target of an x-ray tube for use in apreferred low dosage scanning-beam x-ray imaging system.

FIGS. 3A, 3B, 3C, and 3D are partial cross-sectional representations ofcollimation grids useful in the inventive device.

FIG. 4 is a perspective exploded diagram showing assembly for apreferred embodiment of a collimator grid.

FIG. 5 is a functional representation of components of an x-ray tube fora scanning-beam x-ray imaging system.

FIG. 6 is a diagram showing the target end of a preferred x-ray tube fora low dosage scanning-beam x-ray imaging system.

FIG. 7 is a diagram showing the axes of x-ray beams for a stereoscopicscanning-beam x-ray imaging system.

FIG. 8A depicts a single nonsegmented detector that is smaller in widththan the x-ray beam emitted from an apertured x-ray source.

FIG. 8B depicts an x-ray beam from a single aperture of an aperturedx-ray source interacting with a multi-detector array.

FIG. 8C depicts the axes for x-ray beams from a number of apertures ofan apertured x-ray source passing through the same image pixelinteracting with a segmented detector array.

FIG. 8D is a diagram of x-ray beams from two apertures of an x-raycollimation grid interacting with an object under investigation atvarious distances from the x-ray source.

FIG. 9 is a front view of a 5×5 detector array.

FIG. 10 is a functional representation of one row or column of detectorelements for a 5×5 detector array for an embodiment of a low dosagescanning-beam x-ray imaging system.

FIG. 11 depicts an embodiment of a detector element for a low dosagescanning-beam x-ray imaging system.

FIG. 12 is a diagram of a front view of a 3×3 multi-detector array foran embodiment of a low dosage scanning-beam x-ray imaging system.

FIG. 13 depicts x-ray paths of an x-ray beam emanating from a singlecollimator aperture passing through an object plane to a multi-detectorarray.

FIG. 14 depicts x-ray paths of multiple x-ray sub-beams emanating frommultiple apertures passing through a single pixel to a multi-detectorarray.

FIG. 15 is a front view of the presently preferred embodiment of a 96element scintillator array.

FIG. 16 is a diagram showing a preferred low dosage scanning-beam x-rayimaging system utilizing negative feedback to control x-ray flux.

FIG. 17 is a layout arrangement plan for FIGS. 18-25. FIGS. 18-25provide a single functional block diagram of components of a preferredscanning-beam x-ray imaging system.

FIGS. 18 and 19 are partial functional block diagrams comprising apreferred x-ray source for a scanning-beam imaging system.

FIG. 20 is a partial functional block diagram comprising a preferreddual multi-detector array for a scanning-beam imaging system.

FIGS. 21 and 22 are partial functional block diagrams comprising apreferred monitor for a scanning-beam imaging system.

FIGS. 23 and 24 are partial functional block diagrams comprising apreferred scan generator for an scanning-beam imaging system.

FIG. 25 is a partial functional block diagram depicting the functionalinterconnection of the functional blocks of FIGS. 20, 22 and 23.

FIG. 26 is a partial cross-sectional representation of a preferredreal-time eye assembly.

FIG. 27 is a diagram of a top view of a preferred 96-channelphotomultiplier tube.

FIG. 28 is a partial cross-sectional side view of the photomultipliertube of FIG. 27.

FIG. 29 is a key to FIGS. 29A-D. FIGS. 29A-D are a schematic of apreferred signal conditioning amplifier.

FIG. 30 is circuit diagram of the input and output connectors for apreferred discriminator.

FIG. 31 is a schematic of a preferred discriminator.

FIG. 32 is a key to FIGS. 32A-G. FIGS. 32A-G are is a schematic ofpreferred digital-to-analog converters which provide gain and thresholdcontrol signals to the signal conditioner.

FIG. 33 is a key to FIGS. 33A-I. FIGS. 33A-I are is a circuit diagram ofthe preferred interface connectors between the DACs of FIGS. 32A-G andthe signal conditioning amplifier circuit of FIGS. 29A-D.

FIG. 34 is a circuit diagram of a preferred buffer amplifier forthreshold control signals.

FIG. 35 is a key to FIGS. 35A-D. FIGS. 35A-D are a circuit diagram ofthe preferred connectors between the discriminator of FIG. 31 and thepreferred image reconstruction board.

FIG. 36 diagrammatically depicts a preferred detector arrangement of 144logical detector elements in the presently preferred imagereconstruction engine.

FIG. 37 comprises a partial functional diagram of a string counter in apreferred image reconstruction engine.

FIG. 38 is a diagram showing the preferred beam, alignment octantarrangement.

FIG. 39 comprises a partial functional block diagram of an imagereconstruction engine.

FIG. 40 is a key to FIGS. 40A-E. FIG. 41 is a key to FIGS. 41A-B. FIGS.40A-E and FIGS. 41A-B comprise schematics of the preferred real time eyeoptical to electrical and electrical to optical conversion circuitry.

FIG. 42 is a key to FIGS. 42A-E. FIGS. 42A-E are a schematic of thecontroller for the image reconstruction engine and gain & alignmentcircuitry.

FIG. 43 is a key to FIGS. 43A-B. FIGS. 43A-B are a diagram showing thepreferred input sensor connectors between the photomultiplier tube andthe signal conditioning circuits in the Real-time eye.

FIG 44A is a key to FIGS. 44A-1 to A-4. FIG. 44B is a key to FIGS. 44B-1to B-4. FIGS. FIGS. 44-1 to A-4 and 44B-1 to B-4 comprise schematics ofthe preferred octant counters.

FIG. 45 is a schematic of a preferred gain & alignment ALU.

FIG. 46 is a schematic of a preferred gain & alignment engine.

FIG. 47 is a key to FIGS. 47A-D. FIGS. 47A-D are a schematic of thepreferred memory for the preferred image reconstruction engine and gain& alignment circuitry.

FIG. 48A-I are schematics of string counters for strings one throughnine for a preferred image reconstruction engine.

FIG. 48A is a key to FIGS. 48A-1 to A-2. FIG. 48B is a key to FIGS.485-1 to B-2. FIG. 48C is a key to FIGS. 48C-1 to C-2. FIG. 48D is a keyto FIGS. 48D-1 to D-2. FIG. 48E is a key to FIGS. 48E-1 to E-2. FIG. 48Fis a key to FIGS. 48F-1 to F-2. FIG. 48G is a key to FIGS. 48G-1 to G-2.FIG. 48H is a key to FIGS. 48H-1 to H-2. FIG. 48I is a key to FIGS.48I-1 to I-2.

FIG. 49 is a schematic of a preferred normalization PROM for a preferredimage reconstruction engine.

FIG. 50 is a schematic of the controller for the preferred imagereconstruction engine.

FIG 51 is a key to FIGS. 51A-C. FIGS. 51A-C are a schematic diagram ofthe preferred output FIFOs for the preferred image reconstructionengine.

FIG. 52 is a schematic of the preferred output FIFO controller for thepreferred image reconstruction image.

FIG. 53 is a key to FIGS. 53A-E. FIGS. 53A-E are a circuit diagram ofthe preferred control logic for the detector controller.

FIG. 54A is a key to FIGS. 54A-1 to A-6. FIGS. 54A-1 to 54A-6 are acircuit diagram of the control logic for the preferred tube controller.

FIG. 54B is a key to FIGS. 54B-1 to B-2. FIGS. 54B-1 to 54B-2 are adiagram of the preferred interface circuitry connecting the tubecontroller circuits with the PC bus within the control computer.

FIGS. 55A-E comprise schematics of the preferred beam controllerinterface. FIG. 55A is a key to FIGS. 55A-1 to A-2. FIG. 55C is a key toFIGS. 55C-1 to C-3. FIG. 55D is a key to FIGS. 55D-1 to D-3. FIG. 55E isa key to FIGS. 55E-1 to E-3.

FIGS. 56A-B comprise schematics of the preferred x-deflection driver.FIG. 56 is a key to FIGS. 56A-1 to A-3 and FIGS. 56B-1 to B-3. FIG. 56Bis a key to FIGS. 56B-1 to B-3.

FIGS. 57A-B comprise schematics of the preferred x-step driver. FIG. 57Bis a key to FIGS. 57B-1 to B-5.

FIG. 58 is a cross-sectional view through a two-dimensional array ofregularly-spaced x-ray sources and a two-dimensional array ofregularly-spaced detectors.

FIG. 59 is a diagram showing focal plane location for one embodiment ofa scanning beam x-ray imaging system.

DESCRIPTION OF THE PREFERRED EMBODIMENTS

The following description of the aspects of the present invention isillustrative only and not to be construed as in any way limiting to theinventive concepts disclosed and claimed herein.

System Overview

Turning to FIG. 1, a scanning-beam x-ray imaging system according to apreferred embodiment of the present invention is diagrammed. The x-raysource is preferably a scanning x-ray source 10 preferably comprising apower supply capable of generating approximately -100 kV to -120 kV,which can operate x-ray tube 10 at about -70 kV to -100 kV. At thisvoltage level, x-ray source 10 produces a spectrum of x-rays ranging to100 keV. As used herein, the term 100 keV x-rays refers to thisspectrum. X-ray source 10 includes deflection yoke 20 under the controlof scan generator 30. An electron beam 40 generated within x-ray source10 is scanned across a grounded anode target 50 within x-ray source 10in a predetermined pattern. Hereinafter, for simplicity, anode target 50is referred to as target 50. For example, the predetermined pattern maybe a raster scan pattern, a serpentine (or "S" shaped) pattern, a spiralpattern, a random pattern, a gaussian distribution pattern centered on apredetermined point of the target, or such other pattern as may beuseful to the task at hand. Presently preferred is the serpentine (or"S" shaped) pattern which eliminates the need in a raster scan patternfor horizontal "fly back."

As electron beam 40 strikes target 50 at focal spot 60, a cascade ofx-rays 70 is emitted and travel outside of x-ray source 10 toward theobject 80 to be investigated. To optimize system performance of thepresently preferred embodiment, a cone of x-ray photons should begenerated that will diverge in a manner that will just cover themulti-detector array 110.

This is preferably accomplished by placing a collimating assemblybetween the target 50 of the scanning x-ray source 10 and themulti-detector array 110, and more preferably between the target 50 andthe object to be imaged. The presently preferred collimating assembly isa collimation grid 90, containing a grid of x-ray transmissive apertures140. Collimation grid 90 is designed to permit passage of only thosex-ray pencil beams 100 whose axes lie in a path that directly interceptsmulti-detector array 110. Collimation grid 90 does not move with respectto multi-detector array 110 while the system is in operation. Thus, aselectron beam 40 is scanned across target 50, at any given moment thereis only a single x-ray pencil beam 100 which passes through object 80 tomulti-detector array 110. This preferred result is in contrast to theresult in FIG. 2, which depicts the distribution of x-rays 70 from ascanning-beam x-ray source in the absence of a collimator assembly. Forpurpose of illustration only, the scatter from x-rays 70 which strikemulti-detector array 110 is not shown in FIG. 2.

The output of multi-detector array 110 is Processed and displayed onmonitor 120 as luminance values. Image processing techniques can be usedto produce a computer driven image on an appropriate display orphotographic or other medium.

The embodiment of the inventive system disclosed herein is a lowexposure system in that it typically exposes the cardiology patient at arate of about 0.09 to 0.33 R/min with a 30 frame/sec refresh ratemeasured at the entrance to the patient, which in conventional systemsunder the same conditions would typically be between 2.0 to 2.8 R/min.Whole body exposure with a 30 frame/sec refresh rate with the presentinventive system will be lower than that for conventional systems aswell.

The X-Ray Tube

FIG. 3 depicts a magnified diagrammatic view of a preferred collimationgrid and target structure. Target 50 is preferably comprised of a targetlayer 129 of a material having good vacuum characteristics and theability to withstand high heat and electron bombardment, which is thenformed upon a beryllium target support 53. Aluminum or other relativelyx-ray transparent materials can be used to fabricate the target support53 as well. A preferred construction of the target layer 129 is a firstlayer of niobium 51 approximately 1 micron thick sputter-deposited uponthe beryllium target support 53 to which is then sputter-deposited asecond layer of tantalum 52 approximately 5 microns thick. Thisstructure is presently preferred because niobium has a thermalcoefficient of expansion intermediate to the coefficients of thermalexpansion of tantalum and beryllium, thus reducing or preventingmicrocracking due to thermal cycling of the target as the electron beam40 scans across the target. Another embodiment is a layer of tantalumapproximately 5 microns thick sputter deposited directly on theberyllium target support 53. Yet another embodiment is a layer oftungsten-rhenium approximately 5 microns thick sputter-deposited on theberyllium target support 53. Still another embodiment is a layer oftungsten approximately 5 to 7 microns thick sputter deposited on theberyllium target support 53. Tantalum, tungsten and tungsten-rhenium arepresently preferred for use in target layer 129 because they haverelatively high atomic numbers and densities and readily emit x-rayswhen bombarded by an electron beam. Tungsten's high melting point of3370° C. and good vacuum characteristics make it suitable for the hightemperature and hard vacuum conditions within the x-ray source. Tantalumand tungsten-rhenium have similar characteristics as known to those ofskill in the art. The thicknesses of the target layers are preferablyselected so that they are approximately equivalent to the distancenecessary to efficiently convert 100 keV electrons to x-rays.

Beryllium is presently preferred for target support 53 because it isstrong and does not significantly attenuate or scatter the x-raysemitted from target layer 129. The thickness of beryllium target support53 is preferably about 0.5 cm. In the presently preferred embodiment ofthe present invention, target support 53 should be constructed as thinas possible subject to the physical constraint that it must be strongenough to withstand the pressure gradient of one atmosphere across it.

A cooling chamber 54 is preferably located between the target support 53and collimation grid 90.

Collimation grid 90 preferably consists of an array of apertures 140,the axes of each, according to one preferred embodiment of the presentinvention, are oriented or pointed toward multi-detector array 110. Thatis to say that the axes of apertures within the collimation grid 90 arenot parallel to each other and form an acute angle to the lineperpendicular to the output face 260 of the collimation grid 90. Forexample, a collimation grid for a chest x-ray application may compriseapertures forming an angle with a line perpendicular to the output face260 of the collimation grid 90 of between 0° at the center of thecollimation grid 90 to as much as 20° at the edge of the grid 90. Amammogram application on the other hand may have a collimation grid 90comprising apertures forming an angle with a line perpendicular to theoutput face 260 ranging to 45° at the edge of the grid. Thus, adifferent collimation grid 90 may be selected and inserted for use indifferent medical applications.

The number of apertures 140 in collimation grid 90 may correspond to thenumber of image pixels to be generated on the monitor. For example, 500by 500 to 1024 by 1024. Alternatively, the image pixel to aperture ratiomay be increased, i.e., fewer apertures than image pixels may be used,in conjunction with the technique of "sub-sampling" discussed below. Thesystem spatial resolution may be determined, in part, by the pitch ofthe apertures in collimation grid 90. The precise number of aperturessuggested above is illustrative only, and is not intended in any way tobe limiting.

Some of the factors preferably used to determine the thickness ofcollimation grid 90 and the diameter of apertures 140 are the distanceof the multi-detector array 110 from target 50, which is presentlypreferably 94.5 cm (37.2 in), the desire to significantly attenuate allx-rays 70 not aimed at the multi-detector array 110, and the size of themulti-detector array 110 (not shown in this figure). Apertures 140, asviewed from output face 260, are preferably laid out in a rectangularrow and column pattern containing a substantially circular boundary 25.4cm (10 in) in diameter forming a circular active array. The aperturearray may, however, be of any convenient layout to resolve the image ofobject 80. Further, the electron beam 40 may be scanned in a patternwhich employs only a portion of the apertures 140. The circular activearea according to one preferred embodiment of the present invention hasa diameter of approximately 500 apertures.

The x-ray absorbent portion 150 of preferred collimation grid 90 isdesigned to absorb errant x-rays so that they do not illuminate object80. This is accomplished by fabricating the preferred collimation grid90 with sufficient thickness so that the x-ray radiation passing throughan aperture 140 towards the multi-detector array 110 is substantiallygreater than the cumulative x-ray radiation passing through x-rayabsorbent portion 150 in all directions other than toward multi-detectorarray 110. Such errant x-rays would provide the object 80 and attendingstaff with x-ray dosage but contribute no meaningful information to theimage.

Collimation grid 90, as shown in FIG. 3A, is preferably fabricated froma number of sheets 147 of x-ray absorbing materials having apertures 140therethrough to form an x-ray pencil beam 100 as the x-rays pass throughthe collimation grid to the multi-detector array 110. The material usedfor the sheets 147 can be formed from x-ray opaque materials such asmolybdenum, brass, lead, copper, tungsten, tantalum, gold or any ofthese used in combination. The collimation grid 90 is preferablyfabricated of 50 thin sheets of 0.0254 cm (0.010 in) thick molybdenumwhich are stacked and held together by end plates 258 and 259.Molybdenum is a preferred material for sheets 147 because it readilyabsorbs x-rays so that x-rays generated by x-ray source 10 which are notdirected to multi-detector array 110 will be absorbed before theyimpinge upon object 80, which, of course, may be a human patient. Theend plates 258 and 259 are constructed from an x-ray transmissivematerial, preferably aluminum. Aluminum is a preferred material forplate 259 to minimize x-ray generation in molybdenum collimation sheets147.

Alternatively, collimation grid 90 can be formed of sheets 147fabricated out of both high atomic number materials and low atomicnumber materials to minimize the amount of fluorescent K x-rays whichemanate to the patient. Fluorescent K x-rays, which are generated by theinteraction of the x-rays emanating from the target 50 with thematerials of the collimator sheets 147, are typically undesirablebecause they increase patient exposure without contributing to theformation of an x-ray image. The K x-rays can be attenuated by plate259, but plate 259 should also be transparent to pencil beam 100. Thisis preferably accomplished by using materials of low atomic number forsheets 137, preferably brass, since low atomic number materials producelow energy K x-rays which can be strongly attenuated by plate 259. Forexample, a sheet of aluminum 1 mm thick for plate 259 will typicallyreduce the K x-ray intensity of brass by approximately 99.9%, whilebeing relatively transparent to the higher energy x-rays of pencil beam100. Brass is therefore a superior material for sheets 137 from thepoint of view of stopping K x-rays but, by itself, provides inadequateattenuation for the x-rays emanating from target 50 which do not passthrough collimator aperture 140. Therefore collimator sheets 137preferably comprise a combination of materials, with higher atomicnumber materials such as tungsten, lead or molybdenum at the side ofcollimation grid 90 closest to the target 50 and low atomic numbermaterial such as brass on the side closest to the object 80 to beimaged. Presently preferred is a combination of molybdenum and brass,which provides for high collimation grid is efficiency while producinglow energy K x-rays which are strongly attenuated by plate 259.

The apertures 140 of collimation grid 90 are preferably either round orsquare in cross section. Other shapes could also be used, particularlyhexagons, although the shape of the aperture holes should preferablymatch the shape of the multi-detector array, since the aperture shapeaffects the shape that x-ray beams will tend to diverge into. Forexample, the presently most preferred round aperture hole will tend toproduce an x-ray beam that diverges into a circular shaped beam at themulti-detector array. Therefore, if round apertures are used, themulti-detector array is preferably circular to maximize its exposure andcoverage to the circular x-ray beams.

If square apertures 140 are used they should preferably be 0.0381 cm(0.015 in) by 0.0381 cm in dimension while round apertures arepreferably 0.015 in (0.038 cm) in diameter. Both square and roundapertures yield a cross sectional area at multi-detector 110 that isabout 1/100 the cross sectional area of detectors of known x-rayfluoroscopes. The cross sectional area of the face of the multi-detectorarray 110 is much smaller than in known conventional systems. As aresult, x-rays scattered at the object miss the multi-detector array anddo not tend to fog the image as they do in conventional systems whichtypically utilize relatively large surface area detectors.

The presently preferred method for fabricating the collimation grid 90is by photo-chemical milling or etching. Photo-chemical milling ispresently preferred because it is cost effective and accurate. Accordingto one embodiment of this method, a set of 50 photo masks is created toetch holes or interstices into 50 thin sheets of 0.0254 cm (0.010 in)thick material. In an alternate embodiment, a set of 100 photo masks iscreated to etch holes or interstices into each side of the 50 thinsheets of 0.0254 cm (0.010 in) thick material. The etched sheets arethen preferably stacked, aligned and held together to form a gridassembly having a plurality of stepped apertures, each of apredetermined angular relationship with respect to the sheets. FIG. 3Ashows an embodiment of the preferred collimation grid 90. This variationincludes a number of x-ray absorbing sheets 147 having individualapertures with a constant cross-section (however, the cross-section neednot be constant). The resulting aperture 140 has a steppedconfiguration, as shown while allowing the x-ray pencil beam 100 to passthrough to the multi-detector array 110. The variation shown in FIG. 3Bis quite similar to that shown in FIG. 3A except that the individualapertures formed in x-ray absorbing sheets 146 are themselves stepped.These stepped apertures may be made by milling or chemical etching fromeach side of sheet 146 with a slight offset as described above so as toresult in the configuration shown. The FIG. 3B configuration is highlydesirable because less x-ray energy need be absorbed within the steppedapertures 140 of collimation grid 90 and consequently, the x-ray flux atthe edge of the x-ray beam 100 is not attenuated as much as in thevariation shown in FIG. 3A. X-rays are generally unaffected by theroughness of the channels due to the stepped surface, and even if theyare scattered within the aperture, the scattering will not measurablyaffect the resultant beam. The stepped apertures shown in FIGS. 3A and3B can also be beneficial in controlling the K x-ray intensity asdiscussed in U.S. Pat. No. 2,638,554, issued to Bartow et. al., entitled"Directivity Control of X-rays."

FIG. 4 shows a preferred method for assembling the preferred collimationgrid assembly 90 from etched sheets 91. Preferably 50 etched sheets 91are each provided with alignment holes or alignment apertures 94.Alignment pegs 95 are placed in each alignment aperture 94 to align theetched sheets 91. The assembled sheets 91 and pegs 95 are placed inaluminum ring 359. Aluminum ring 359 is provided with a vacuum port 370which may, after assembly, be sealed with pinch off 375. Aluminum sheet365 which is preferably 0.1 cm in thickness is bonded and sealed with avacuum adhesive to upper surface 380 of ring 359. Aluminum sheet 360 issimilarly bonded to a lower surface 385 of ring 359. A partial vacuum isthen pulled through port 370 and the port 370 is then sealed at pinchoff 375. The partial vacuum causes relatively x-ray transparent aluminumsheets 360 and 365 to provide a clamping action tending to hold etchedsheets 91 together and in alignment to form a collimation grid 90. Thepresently preferred tolerance for the aperture center-to-center distanceis ±0.00127 cm (0.0005 in) without cumulative error. The presentlypreferred tolerance on the aperture sizes is ±0.00254 cm (0.001 in). Forease in assembly the diameters of apertures in every other sheet may befabricated to be larger than the diameter at the output face of thecollimator. Thus, only every other sheet need be carefully aligned. Thematerial used for the sheets 91 as discussed above can be molybdenum,brass, lead, copper, tungsten, tantalum, gold or any of these used incombination. Molybdenum is a preferred material for use in the sheets91, but more preferred at present is a combination of molybdenum andbrass.

In an alternate method to fabricate collimator grid 90, the alignmentholes 94 are etched along with the apertures 140. However, due to thedifferential between the sizes of the apertures 140 and the alignmentholes 94, undercutting of the apertures 140 occurs because the time foretching is governed by the time it takes to etch the large holes. In apresently preferred alternative method, pilot alignment holes smallerthan the final size of the alignment pegs are etched into the sheets 91.The etched sheets 91 then undergo an additional procedure such asreaming to enlarge the pilot alignment holes to the desired diameter.The finished sheets are aligned and clamped together as previouslydescribed.

Alternative methods for fabricating collimation grid 90 include electronbeam machining, drilling, mini-machining, and laser drilling. Drillingand laser drilling are useful for generating round holes, but a drawbackis the relative difficulty in generating square holes with eithermethod. In addition, these non-etch methods typically require greatertime and costs when compared to the above described etching methods.

More details of the preferred scanning x-ray source 10 are shown inFIGS. 5 and 6. Grid controlled electron gun 185 is preferably locatedopposite the face of x-ray tube 10 and is operated at a potentialbetween -70 kV to -120 kV. The electron beam 40 emanating from electrongun 185 can be controlled in amplitude and can be rapidly reduced tozero by the application of an appropriate voltage to a control grid 182.Grounded target 50 is preferably located at the face of the tube andelectron beam 40 is preferably emitted from electron gun 185 towardstarget 50. A grounded anode 186 is preferably located near electron gun185 and includes an aperture 187 at its center for electron beam 40 toaccelerate the electrons as they pass through. Divergent electron beam40 is accelerated towards anode 186 and passes through aperture 187.Magnetic focus lens 188, preferably of fixed power, causes the electronbeam 40 to become convergent so that it strikes target 50 at focal spot60. Focal spot 60 preferably has a diameter of 0.3 mm. Varying currentsflowing in the coils of magnetic deflection yoke 20 preferably deflectthe electron beam 40 so that focal spot 60 moves over the surface oftarget 50 in the previously mentioned preferred serpentine pattern.Dynamic focus coil 183 is energized by a current which varies insynchronism with the varying current in deflection yoke 20 to maintainthe preferred 0.3 mm diameter for focal spot 60 as the electron beam 40is scanned over the surface of target 50. The tube is preferablyfabricated to have a 25.4 cm (10 in) diameter sweep area to correspondwith the circular active area of the collimator grid 90. Electron beam40 intersects target 50 at an angle of up to about 30° at theextremities of the circular active area. When the x-ray source 10 is inuse, no more than one aperture 140 (possibly two for stereo) ofcollimation grid 90 will be passing an x-ray pencil beam 100 at anygiven instant. According to one preferred embodiment, the electron beam40 may be shut off by application of a short rise voltage pulse tocontrol grid 182 when focal spot 60 is not positioned directly in frontof an aperture 140. Thus the x-ray tube may be operated effectively in ascanned-pulsed mode to reduce power consumption approximately 25% andheating of the target 50.

Turning to FIG. 6, a cross-sectional view of the front portion of thepreferred x-ray source 10 is depicted. The interior of the x-ray source10 is maintained at a vacuum. Target 50 as discussed above is comprisedof a suitable target material deposited on beryllium target support 53which is 0.5 cm thick. Forward of beryllium target support 53 is coolingjacket 54 which is preferably 0.2 cm thick and may be adapted to carrywater, forced air or preferably Fluorinert. Aluminum grid supports 360and 365 are each preferably 0.1 cm thick and help support collimationgrid 90 which is preferably 1.27 cm (0.5 in) thick. Aluminum gridsupports 360 and 365 together with the beryllium target support 53 andthe coolant in cooling jacket 54 preferably form an x-ray filter whichfilters out low energy x-rays. The presently preferred x-ray source isdescribed more fully in copending U.S. patent application Ser. No.08/386,884, which has been incorporated herein by reference in itsentirety.

Stereoscopic X-ray Imaging

Turning now to FIG. 7, a collimation grid having more than one focalpoint may be provided so that stereoscopic x-ray images may be obtained.If, for example, the axes 101 of the x-ray pencil beams 100,corresponding to the aperture axes of every other row of apertures 140ain grid 90 are pointed at focal point F1 at the center of multi-detectorarray 92 and the aperture axes of the remaining apertures are pointed atfocal point F2 at the center of multi-detector array 93, one can scanthe apertures in a raster or serpentine pattern and create a "line" ofdata from the first multi-detector array, and a line of data from thesecond multi-detector array. Repeating this, it is possible to build uptwo complete images, as seen from two distinct angles and therebydisplay them with conventional stereoscopic imaging display systems toprovide a stereoscopic x-ray image.

FIG. 3C depicts how one may construct such a stereoscopic collimationgrid out of layers 147 of x-ray absorptive material. In this embodimentapertures 140A, 140B may diverge from a common first aperture 140 toform a "V" as shown providing separate paths along the "legs" of the "V"for x-ray pencil beams 100A, 100B. There is no requirement, however,that apertures 140A, 140B diverge from a common aperture as shown, butan advantage of the "V"-shaped aperture where the x-rays enter at thecommon aperture or apex of the "V' is that both multi-detector arrays 92and 93 will be illuminated simultaneously, the "V" acting as an x-raysplitter with some of the x-rays going to multi-detector array 92 andsome to multi-detector array 93. This decreases by 50% the powerrequired for the beam current.

Additionally, the system may be designed to such that the commonadjacent holes in the first sheet share common holes in the last sheetlast two sheets as shown in FIG. 3D.

The Multi-Detector Array

To achieve resolutions of several line pairs per millimeter or more atthe object plane, as are required in some medical applications, thespatial resolution limit in known reverse-geometry systems is in largepart determined by the size of the single nonsegmented detector.Generally speaking, a small non-segmented detector can provide highspatial resolution while a large non-segmented detector provides highcollection efficiency. It has in part been this trade-off that has beena problem in developing low dosage x-ray imaging systems. Other partshave been the inability to fabricate a suitable collimator and the lackof a high efficiency x-ray scintillator also having a fast decay time.

When such a detector is small to increase resolution, a large proportionof the x-rays emitted by target 50 are unused by the single detector250, as shown in FIG. 8A even when a collimator grid 90 is used. Thisis, in fact, how industrial reverse-geometry scanning-beam x-rayinspection systems are designed, where dose is usually not aconsideration. Accordingly, while one can decrease the size of adetector by placing, for example, a lead washer in front of the singledetector 250 and thereby increase spatial resolution, the x-rayintensity and/or exposure time would have to be increased to maintaincontrast resolution.

By fabricating a multi-detector array having a large area subdividedinto multiple smaller detector array elements (e.g., as shown by thefront view of the multi-detector array 110 in FIG. 9) a large capturearea is achieved, while simultaneously through image reconstructiontechniques described herein retaining an image resolution that iscomparable to the size of a single small detector element withoutincreasing x-ray intensity an/or exposure time.

The resolution defined by the individual detector elements 160 ismaintained by distributing and summing the outputs from the individualdetector elements 160 into a memory buffer in which each address, i.e.,image pixel, corresponds to a specific location in the object plane 280.As an electron beam 40 is moved discretely across the target 50,illuminating the area behind selected apertures 140 of the collimationgrid 90, the address, to which the output of a given individual detectorelement 160 is added, changes. The imaging geometry is shown in FIG. 8Band 8C. In FIG. 8B a single x-ray beam 100 is shown along with how itgenerates information for 5 image pixels. Effectively, the single x-raypencil beam 100 emanating from individual aperture 141 is divided intox-ray micro-beams, the number of x-ray micro-beams created correspondingto the number of individual detector elements 160 which comprise themulti-detector array 110. In the case shown in FIG. 8B. The axes of fivex-ray micro-beams 141a, 141b, 141c, 141d and 141e are shown. In FIG. 8Cthe sequential positions of the axes of the x-ray micro-beams from x-raypencil beams 100 emanating from five consecutive apertures 141 through145 illuminating a single image pixel ("IP") are shown. The outputs fromthe five individual detector elements 161, 162, 163, 164 and 165receiving the x-ray flux from the five x-ray micro-beams, 145a, 144b,143c, 142d and 141e respectfully, are added together to provide theluminance for the single pixel IP.

Stated differently, the output for each of the individual detectorelements 160 is stored for later summation in an image buffer, at amemory address that corresponds to a very small specific region in theobject plane 280, e.g., a single image pixel.

Accordingly, in one embodiment the memory storage address for the outputof each individual detector element 160 changes with the position of thescanning x-ray beam 40 in an ordered fashion such that each memoryaddress contains the sum of the radiation passing through a specificimage pixel or spot in the object plane 280. In this way the spatialresolution of the system is determined by the size of a singleindividual detector element 160, while the contrast resolution of thesystem is determined by the area of all of the individual detectorelements comprising the multi-detector array 110.

An additional benefit of this multi-detector array imaging geometry isthat the depth of field of the object plane 280 is narrowly defined.Structures lying in front of or behind it will be blurred (out offocus). X-ray pencil beams from a first aperture 141 and a secondaperture 142 are depicted in FIG. 8D passing through an object plane 280a distance S₀ from apertures 141,142 and passing through a plane 281 adistance S₁ from apertures 141, 142 where S₁ >S₀. The bubbles representimage pixels IP₁ through IP₁₀. As can be readily seen, the resolution atS₁ is less than that available at S₀. This feature provides for improvedlocalization and visualization of detailed structures in the plane ofinterest 280, while providing an adequate depth of field that may bemodified by the system geometry.

The multi-detector array 110 of the presently preferred embodimentcomprises 96 individual detector elements 160 arranged in a pseudo-roundarray of square scintillator elements 0.135 cm on a side disposed withina circle of diameter about 1.93 cm (0.72 in). This number of individualdetector elements is merely illustrative. The preferred multi-detectorarray 110 is described more fully in co-pending patent application Ser.No. 08/387,292, which has been incorporated herein by reference in itsentirety.

The Detector Elements

Conventional image intensifier technology typically has basicconstraints that limit a system's sensitivity. One of the objects of thepresent invention is to provide a scanning-beam x-ray imaging systemwhich will result in the subject under examination being exposed to thelowest possible level of x-rays commensurate with achieving imagequality adequate to meet the requirements of the procedure beingperformed. This means that the system used to detect the x-ray photonsemerging from the subject preferably has the highest possible detectivequantum efficiency. To achieve this, the scintillating material used inthe individual detector elements preferably has a length in thedirection in which the x-ray photons travel that is sufficient to ensurethat no x-ray photons emerge from the end opposite the incident x-rays,i.e., the x-ray photon energy should be adequately dissipated in thematerial to maximize the output of the detector.

There are several types of individual detector elements which can beused in the presently described scanning-beam x-ray imaging system. Thatwhich is currently preferred comprises a scintillator in which x-rayphoton energy is converted to visible light energy and the lightintensity is then converted to an electrical signal by means of aphotomultiplier, photo diode, CCD or similar device. Because theinformation from each aperture must be obtained in a very short timeperiod, the scintillating material should have a fast response and aminimum afterglow time. Afterglow is the phenomenon wherein thescintillator continues to emit light after the stimulating incidentx-rays have ceased. Even faster response and shorter afterglow times arerequired if x-ray intensity measurements are obtained using thepreferred x-ray photon counting technique.

Plastic scintillators, such as organic loaded polystyrene, are suitablefrom a standpoint of speed in that they have the required fast responseand minimum afterglow characteristics. However, plastic scintillatorshave a relatively small x-ray photon interaction cross section so thattheir linear x-ray absorption coefficients are also small in value. Theconsequence is that a considerable thickness is required to absorb x-rayphotons. For 100 kV x-rays, a typical plastic scintillator should beabout 28 cm (1.1 in) thick to capture 99% of the incident x-rays. Morepreferred materials at present (and in order of preference) are: (1) YSO(cerium doped yttrium oxy-orthosilicate) available from Airtron (Litton)of Charlotte, N.C.); (2) LSO (cerium doped lutetium oxyorthosilicate)available from Schlumberger, Inc.); and (3) BGO (bismuth germanate,available from Rexon Components, Inc. of Beachwood, Ohio). YSO and LSOare advantageous in that they may be used at room temperature. BGO mustbe heated to about 100° C. in order to achieve a suitable light outputdecay period of the order of 50 nanoseconds. These scintillatingmaterials need not be as long as the plastic scintillator and aretypically effective at a length of 0.10 cm, and preferably at a lengthof several millimeters.

According to one embodiment of the present invention, multi-detectorarray 110 comprises at its input face a pseudo-round array of 96 denselypacked scintillators including two rows of 12 and two columns of 12 atits horizontal and vertical midplanes spaced a distance of preferably91.4 cm (36 in), and more preferably 94.5 cm (37.2 in), from the x-raysource 50. (FIG. 15) A square 5 by 5 and a square 3 by 3 array are alsocontemplated as is a non-square array of scintillators square crosssections filling a circle about the center of the multi-detector array.If scintillators 170 have parallel sides, x-rays entering near the edgesmay strike the scintillator walls and pass through to a neighboringscintillator of an adjacent detector element causing the detectedelement to generate an output seemingly from the wrong spatial positionin the subject with consequent degradation of the image quality. Thisproblem is addressed by placing shielding material between theneighboring scintillators. While some x-ray photons are lost, in thatthey may not generate a light photon, due to the image reconstructionmethod employed, the resultant image is not affected to any substantialdegree.

Referring now to FIG. 10, according to one presently preferredembodiment of the detector element 160, each scintillator element 170 ispreferably in contact with a light pipe or fiber-optic coupler 180 whichoptically couples each scintillator element 170 with a correspondingphotomultiplier tube 190 or solid state detector. A coupling oil ispreferably used between each end of the fiber optic coupler to ensureproper transmission at the interfaces. Alternatively, scintillators 170may be located in close physical proximity to their correspondingphotodetectors, eliminating the fiber optic coupler.

FIG. 11 shows an alternative configuration of a detector element 160without the optical coupler. An x-ray opaque sheet 200 with apertures210 corresponding to each detector element 160 is disposed in front ofmulti-detector array 110. Each detector element 160 is enclosed in alight tight enclosure 220 which may also be x-ray opaque. A lightblocking window 230, preferably made of thin aluminum sheet is locatedat the front of light tight enclosure 220. Light blocking window 230 isx-ray transmissive. Within light tight enclosure 220 is a scintillatorelement 170 in close proximity to a photomultiplier tube 190 which ispreferably electrically connected to a pre-amplifier 240.

Alternatively, scintillators could be placed in direct or close contactwith an array of photo diodes, photo transistors or charge coupleddevices (CCDs) to achieve a rugged and compact detector. Where solidstate devices, particularly CCDs, are used, cooling, such as with aPeltier-type cooler, or the like, may be employed to increase thesignal-to-noise ratio of the device. Alternatively, the scintillatorarray could be placed in direct or close contact with one or moreposition sensitive photomultiplier tubes which provide an output signalwhich identifies the position coordinates of the light source as well asits amplitude.

According to a presently preferred embodiment of the present invention,the scintillators are coated along their lengths and the input face witha material which reflects light, such as preferably titanium dioxide, toprevent light from escaping from or entering the scintillators and toaid in internal reflection within the scintillators.

According to a preferred embodiment of the present invention, eachscintillator element 170 is isolated from its adjacent scintillatorelements 170 by a thin sheet 171 of a highly x-ray opaque material suchas, for example, gold or lead. Sheets 171 may be about 0.0102 cm (0.004in) to 0.0127 cm (0.005 in) thick and is most preferably 0.0051 cm(0.002 in) to 0.0127 cm (0.005 in). An example of the position of sheets171 between the scintillators 170 is shown in FIG. 12.

The area of the circular active area of collimation grid 90 ispreferably larger than the area of multi-detector array 110. Thus theaxes of the x-ray pencil beams 100 emitted from the respective apertures140 of collimation grid 90 all converge toward the multi-detector array110 while each individual x-ray pencil beam 100 diverges, or spreads, aswould a flashlight beam to cover the face of the multi-detector array110.

Image Reconstruction

An important advancement of the present invention concerns theapplication of an image reconstruction system to obtain high qualityx-ray images. The output of the multi-detector array is Preferably notapplied directly to the luminance input of a video monitor. Instead,digitized intensity data for each image pixel are stored in a discreteaddress in a "frame store buffer". More than one such buffer may be usedin certain applications. Pixel addresses within the buffer can berandomly accessed and the intensity value can be manipulatedmathematically. This function has application in applying various imageenhancement algorithms and it allows for pixel assignment of the datafrom discrete segments of the detector array.

Referring to FIG. 13, this diagram illustrates the divergence of asingle x-ray pencil beam 100 from aperture 140 to the multi-detectorarray and how it intersects an object 80 (not shown) at object plane280. Image pixel 121 is just one of the image pixels comprising thex-ray pencil beam intersection area 122 of object plane 280. Arepresentative sample of the axes 102 of the x-ray micro-beams createdby having a segmented array are also shown. In FIG. 13, x-ray pencilbeam 100 is shown emitted through a single aperture 140 of collimatorgrid 90. X-ray pencil beam 100 as it exits aperture 40 will divergeforming a cone having a cross section the size of the aperture as itexits the aperture to a cross section covering the scintillators of thedetector elements of the multi-detector array by the time it reaches the96 element multi-detector array 110. The 96 element multi-detector array110 is preferably positioned and designed such that the area of the coneof the x-ray beam 100 just covers the surface area of the multi-detectorarray 160 when the x-ray pencil beam 100 intersects the face of themulti-detector array.

As x-ray pencil beam 100 passes through an object 80, information aboutobject 80 will be detected by the multi-detector array 110 as x-rayintensity values. Because multi-detector array 110 is composed of 96separate detector elements, each detector element 160 will detect onlythe intensity value for the particular x-ray micro-beam 101 of a segmentof x-ray pencil beam 100 that it intersects with. The cross sectionalshape and area of the x-ray micro-beams will correspond to the crosssectional area and shape as the input face of the detector elements. Forexample, if the input faces are square, the x-ray micro-beam will have asquare cross section. The x-ray pencil beam 100 emitted from eachaperture 140 on collimator grid 90 will therefore generate one group of96 separate or discrete pieces of information (the intensity value ateach detector element) about 95 areas of object 80 in the x-ray pencilbeam's 100 path 122. The intensity information from each of the x-raymicro-beams provide partial image pixel information which can be used tocompile complete image pixel information for each image pixel in adesired plane of object 80.

FIG. 14 illustrates the axes 102 of all of the x-ray micro-beams fromall of the apertures 140 that intersect a single image pixel 121 inobject plane 280 as they travel to the multi-detector array 110. Thisimage pixel group of x-ray micro-beams is ultimately processed togenerate an image pixel on a video monitor. In a preferred embodiment ofthe scanning-beam x-ray system, the apertures 140 on collimator grid 90will generate x-ray pencil beams 100 in a predetermined pattern. Asx-ray pencil beams 100 pass through an object, x-ray micro-beams 101from adjacent and nearby apertures will intersect at, for example, point121 (e.g. an image pixel) in the object. The intensity of each of thesex-ray micro-beams 101 from these x-ray pencil beams 100 after they passthrough the object provide information about these intersecting pointsin the object. In this preferred embodiment, each intersecting point onthe object can therefore be considered a single-image "pixel" 121. Inaccordance with the techniques explained in more detail herein, eachimage pixel 121 is preferably mathematically reconstructed from theintensity information of the separate x-ray micro-beams 101 that weregenerated by the detector elements 160 for each of the emitted x-raypencil beams 100 from, for example, the image pixel group of aperturesthat generated x-ray micro-beams whose axes passed through the object atthat point, image pixel 121.

According to a preferred embodiment of the present invention, the outputimage would preferably consist of up to about 250,000 pixels, arrangedin 500 rows and 500 columns. For the purpose of the explanatory examplebelow, it is assumed that the scanning x-ray source is momentarilycentered upon the point, P, located at row 100 and column 100 ofcollimation grid 90. It is further assumed in regard to this embodimentthat the detector array 110 consists of a square 3-by-3 multi-detectorarray 110 containing 9 detectors 179 (FIG. 12) and that each detector179 is sized so as to intercept all of the x-ray emissions associatedwith a single image pixel. Other array configurations obviously may beused as are detailed herein.

The resulting intensity values, preferably digitized from the individualdetectors of the multi-detector array 110, may be assigned to pixelbuffer addresses as follows:

detector 1--row 99, column 99

detector 2--row 99, column 100

detector 3--row 99, column 101

detector 4--row 100, column 99

detector 5--row 100, column 100

detector 6--row 100, column 101

detector 7--row 101, column 99

detector 8--row 101, column 100

detector 9--row 101, column 101

In this example, a corresponding pattern of data assignment is repeatedas the scanning x-ray beam passes behind all of the pixels.

In the displayed image, with a sub-sampling ratio of 1:1, the numericalvalue of each image pixel is equal to the sum of "n" parts where "n" isthe number of detectors 179 in the multi-detector array 110 (in thisexample, n=9).

When constructed as shown in this example, the multi-detector array 110together with the image reconstruction method selected, has the effectof fixing the working distance at which optimum focus is obtained andproviding a plane of optimum focus not available in prior artnon-segmented detector scanning-beam imaging systems.

The following parameters must be taken into consideration in design ofthe multi-detector array:

1. The size and shape of the collimated beam from the x-ray source(target 50);

2. The distance between the target 50 and the multi-detector array 110,"SD" (FIG. 8);

3. The distance between the source 50 and the center of the object ofinterest 280, "SO";

4. The desired resolution, or pixel size at the object of interest 80;and,

5. In medical applications, the total area of the multi-detector array110 must be large enough to intercept all of the x-rays in x-ray pencilbeam 100 emanating from the collimation grid 90, to avoid exposing thepatient to x-ray radiation which does not contribute to the image.

In a preferred embodiment of the invention, the distance between thetarget 50 and the exit side 260 of collimation grid 90 is about 2.271 cm(0.894 in), and more preferably 2.54 cm (1.00 in) (see FIGS. 3 and 6).Apertures 140 are preferably round with a diameter of 0.0381 cm (0.015in). If the apertures are square they are preferably 0.0381 cm on aside. The spot size of electron beam 40 on target 50 is preferably about0.0254 cm (0.010 in) in diameter. The multi-detector array 110 ispreferably about 91.4 cm (36 in), and more preferably about 94.5 cm(37.2 in), from target 50. The preferred beam divergence angle of x-raypencil beam 100 is calculated by 2*ARCTAN((spot diameter/2)/((aperturewidth/2)+(spot diameter/2))* (spot diameter). The projected x-ray pencilbeam 100 diameter is SD* TAN (divergence angle). It has been determinedthat the preferred size of the multi-detector array 110 should be about2.54 cm (1 in) in diameter for the more preferred embodiment.

For example, with a multi-detector array size of 2.54 cm (1 in) square,if the object plane to be imaged is 22.86 cm (9 in) from target 50 andthe desired image pixel size is 0.0508 cm (0.020 in) at the objectplane, and the distance from the target 50 to the multi-detector arrayis 91.4 cm (36 in), the projected size of image pixels at the detectorplane 270 is (SD/SO)*pixel size at object, or 0.2032 cm (0.080 in). Thedesired resolution may be obtained by dividing 2.54 cm (1 in) by 0.2032cm (0.080 in) yielding a square multi-detector array having 12 to 13detector elements on a side. Other configurations are possible dependingupon the circumstances in which the invention is to be used.

Outside of the plane of optimum spatial resolution, SO (280 in FIG. 8Dand FIG. 14), spatial resolution will degrade. In some applications,such as imaging of the human heart, degraded spatial resolution outsideof the depth of field of the system may be seen as being advantageousbecause blurring of detail outside of the area of interest may tend toincrease the perception of details within the area of interest.

A number of methods can be used to obtain a useable image from the dataobtained as described above. As described above, a simple convolutionmethod may be used. Two additional methods are presently preferred forobtaining maximal resolution and sensitivity from the captured data.These are called the multi-image convolution method and the multi-outputconvolution method. For both cases, the following is assumed:

Assume there are AP_(x) rows of apertures and AP_(y) columns ofapertures in collimation grid 90 (FIG. 14). Each intersection of acolumn and row is an "aperture point." Those aperture points outside ofthe circular active area of collimation grid 90 are treated as if theycontribute no measured intensity to the image, i.e., they are treated asif they are "dark". Aperture points not illuminated by electron beam 40during a scan are similarly treated as if they contribute no measuredintensity to the image, i.e., they are also treated as if they are"dark".

Turning now to FIGS. 14 and 15, in multi-detector array 110 assume thereare a maximum of DET_(x) rows of detector elements 160 and a maximum ofDET_(Y) columns of detector elements 160 in a pseudo-round arrangement.

ZRATIO is a real number between 0 and 1. If ZRATIO=1, the focus will beat the detector plane 270 and if ZRATIO=0, the focus will be at thetarget 50 plane. If ZRATIO=0.5, then the focus is half way between thetarget 50 and the detector plane 270, and so on. PIXELRATIO is thenumber of image pixels per physical distance between adjacent detectorsin a column or row. For example, if the spacing between image pixelcenters at object plane 280 is 0.01 cm, and the spacing between thecenters of the detector elements at detector plane 270 is 0.1 cm, thenPIXELRATIO=10. FOCUS=ZRATIO*PIXELRATIO. Focus is typically in the rangeof 0.5 to 2.0 and is usually 1.0.

IMAGE is a data array of size DET_(x) by DET_(y) containing theintensity information for a particular scan and corresponding to aparticular aperture point. PIXEL is a 4-dimensional array of sizeDET_(x) ×DET_(y) ×AP_(x) ×AP_(y) which contains the IMAGE data arraysobtained by scanning all (or part of) the apertures. PIXEL is refreshedafter each scan according to one preferred embodiment of the presentinvention.

As the electron beam 40 is scanned across the target 50 surface, it is,in effect, positioned at the center of selected apertures 140, "fired,"and then repositioned. Thus for each firing or pulse an IMAGE array ofdata will be acquired. While these arrays of data could be constructedinto a displayable image having some use directly, more resolution andsensitivity at lower dosage is obtained by combining them.

The first preferred method for combining the images is called themulti-image convolution method. In the multi-image convolution method,an OUTIMAGE array of intensities of size AP_(x) ×AP_(y), which can bedisplayed on a CRT or like display means, is formed by assigning toOUTIMAGE(y,x) the value of: ##EQU1##

The second presently preferred method for combining the AP_(x) ×AP_(y)number of IMAGE data arrays into a useful picture is called themulti-output convolution method. In this case, with a sensor array ofDET_(x) ×DET_(y) sensors there will be DET_(x) ×DET_(y) digitizers (ortheir equivalents, multiplexed) and the same number of pixel summingcircuits. The digitized values from each sensor are called SENSOR(j,i).The final OUTIMAGE array is computed as follows--for each pixel in theoutput image array OUTIMAGE(y,x) for y=1 to AP_(y) and x=1 to AP_(x) !one pixel from each of the DET_(x) ×DET_(y) source images SENSOR(j,i) issummed for j=1 to DET_(y) and ii to DET_(x) ! into destination imagepixel OUTIMAGE(y-j*FOCUS,x-i*FOCUS). Normalization is then carried outover the OUTIMAGE array by dividing each element thereof by DET_(x)*DET_(y).

A further improvement upon these techniques may be obtained byperforming linear interpolation based upon the fractional part of theFOCUS factor.

An advantage of the multi-image convolution method over the multi-outputconvolution method is that the former allows the plane of optimum focusto be selected in software after the data is captured while the latterdoes not. The latter method, however, may be performed quicker wheretiming is a limitation.

Three Dimensional Image Reconstruction

The scanning-beam imaging system described herein may be used togenerate a set of sequential planar images which can then be used toform a tomograph or a three dimensional display of the object 80. Animage set can be analyzed to produce a 3D image consisting of a seriesof images at various depths by re-analyzing the data set with variousvalues of FOCUS. The natural FOCUS values to use are n/DET_(x) orn/DET_(y) where n is an integer from 0 to DET_(x) or DET_(y),respectively. Normally, only the focus values corresponding to planes ofinterest in the object 80 would be analyzed. For example, in thescanning-beam imaging system described in TABLE I (below), the planes offocus would be spaced at approximately 2.54 cm (1 in) intervals aroundthe normal object plane of 22.86 cm (9 in) (plane of optimum focus).

The following preferred formula shows where the sequential planar imagesare located in terms of distance from the target 50. ##EQU2## WhereF_(t) (FOCUS)=Distance from the target to the particular focal plane ofinterest

F_(d) =Distance from the detector to focal plane (distance from thetarget to the detector less F₁)

λ_(s) =Center-to-center spacing between adjacent collimation gridapertures

λ_(d) =Spacing between centers of adjacent detectors 160 within detectorarray 110.

When using the sub-sampling method, the basic method of computation doesnot change--only the data from the collimation grid apertures which arenot "skipped" need be processed. Note that λ_(s) remains the same evenif intervening collimation grid holes are skipped.

Negative Feedback X-Ray Flux Control

Turning now to FIG. 16 an embodiment of the preferred scanning-beamimaging system employing a negative feedback path 305 to control thex-ray flux of x-ray pencil beams 100 is depicted. Negative feedback fromthe multi-detector array 110 can be utilized to control x-ray flux sothat the multi-detector array 110 always sees approximately the sameflux level. In this way when soft tissue (which is relativelytransparent to x-rays) is being scanned, the x-ray flux can be lowered,without degrading the image reducing the overall dosage to the patient(or object). Improved contrast and dynamic range are provided by usingnegative feedback flux control. According to this embodiment,differential amplifier 310 has an adjustable reference level 320 whichmay be set by the user. Negative feedback loop 305 feeds back to x-raytube 10 to control the x-ray flux.

Time Domain Scanning Mode

A time domain x-ray imaging system may also be implemented using theprinciples disclosed herein. In such a system, the time to reach apredetermined measured x-ray flux from the various pixels could becomputed and mapped. Negative feedback control could then be employed toturn off or reduce x-ray flux from apertures corresponding to pixelswhich had reached the predetermined flux level for the scan period inquestion. In this case, the information gathered would be time to fluxlevel and the mapped or imaged information would correspond to timerather than intensity. Such a system has the potential to provide muchhigher signal-to-noise ratios, improved contrast, drastically reducedx-ray dosage to the object under investigation, and improved dynamicrange.

Multiple Energy X-ray Imaging Mode

According to one preferred embodiment of the present invention, two ormore groups of x-ray pencil beams 100 are directed toward one or moredetector arrays through two or more groups of aperture 40. A first groupof x-ray pencil beams 100 has a first characteristic x-ray energyspectrum. A second group of x-ray pencil beams 100 has a differentsecond characteristic x-ray energy spectrum. By comparing the measuredtransmissivities of the first and second group of x-ray pencil beams100, the presence of certain materials in the object under investigationmay be detected. The basic concept of use of differential x-ray imagingin known imaging systems is disclosed, for example, in U.S. Pat. No.5,185,773 entitled "Method and Apparatus for Nondestructive SelectiveDetermination of a Metal" which is hereby incorporated herein byreference.

The two groups of x-rays may be generated in a number of ways. One suchway is by fabrication of a special target 50 having a first material orfirst thickness of a material adjacent to the apertures of the firstgroup of apertures and a second material or second thickness of materialadjacent to the apertures of the second group of apertures. In thismanner, the apertures associated with the first group will preferablyemit x-rays having a first characteristic energy spectrum and theapertures associated with the second group will preferably emit x-rayshaving a second characteristic energy spectrum.

Alternatively, K-filtering (or K-edge filtering) can be used by placingfilter material (such as, for example, molybdenum) within a portion ofthe apertures 140 to produce a similar effect. In this case, a firstgroup of apertures would comprise a first filter inserted therein and asecond group of apertures would comprise a second filter insertedtherein. The second filter could be no filter at all. As in the previouscase, two groups of x-rays having different characteristic energyspectra would be associated with the two groups of apertures.

Once at least two groups of apertures are associated with differentcharacteristic x-ray spectra, it is now easier to detectmicro-calcification (associated with breast cancer and coronary arterydiseases) and other abnormalities not normally visible with broadbandx-rays. For example, by performing a scan of the first group ofapertures to form a first image, then performing a scan of the secondgroup of apertures to form a second image, and dividing the images tohighlight their ratios, it is easier to detect micro-calcification andother such abnormalities with a low dosage scanning-beam x-ray imagingsystem--in real time. Similarly, a multiple detector array arrangementcould be used with group 1 apertures directed toward a first detectorarray and group 2 apertures directed toward a second detector array,etc.

Another embodiment of multiple energy imaging uses an x-ray photoncounting detector system in which the amplitude of the electrical pulsefrom a detected x-ray photon is proportional to the energy (keV) of thephoton and the pulses coming from photons in two or more energy bandsare counted separately. The pulses are separated by amplitude (i.e.,x-ray photon energy) and then counted and processed separately, formingtwo or more separate images. Those images can be displayed as ratios. Itis also possible to rapidly change the selected energy levels todistinguish different density regions in the object. The advantage ofthis embodiment is that it is more flexible than those described above,and does not require special collimation grids, target materials or dualdetectors.

While a number of preferred embodiments have been discussed above forvarious configurations of the present invention, the followingspecifications are illustrative of a presently preferred scanning-beamimaging system according to the present invention:

                  TABLE I    ______________________________________    Grid    Shape:           round    Diameter:        25.4 cm (10 in)    Aperture Pitch:  0.0508 cm (0.020 in), 0.152 cm                     (0.060 in for sub-sampling)    Max. No. of Apertures                     500 (166 for sub-sampling)    in a row or col:    Cir. Active Area of grid:                     506.45 sq. cm (78.5 sq. in)    Number of apertures:                     196,350 approx. (approx.                     21,630 for sub-sampling)    Aperture cross-section:                     round    Aperture Diameter:                     0.0381 cm (0.015 in)    Space between apertures:                     0.0127 cm (0.005 in) (0.045 in                     for sub-sampling)    Grid Output Face to    target surface Dist:                     2.54 cm (1.0 in)    General System Config.    Source-Detector Distance:                     94.5 cm (37.2 in)    Location of Object Plane:                     23.6 cm (9.3 in) from Source    Scan Frequency:  Adjustable to 30 Hz (Unless                     region of interest scanned)    Operating voltage on x-                     70-120 kV    ray tube:    Detector Array    Overall shape:   pseudo-round (per FIG. 15)    Shape of input face of                     square    detector elements:    Size of input face of                     0.142 cm × 0.142 cm    detector elements:    Number of detector                     96    elements:    Array diameter:  1.83 cm (0.72 in)    Field of view at Plane of                     19.05 cm (7.5 in)    Optimum Focus:    Pixel size at Plane of                     0.038 cm    Optimum Focus:    Detector spacing:                     0.152 cm    Resolution:      approximately 20 line pairs/cm    ______________________________________

Accordingly, one embodiment of the scanning-beam imaging systemutilizing a multi-detector array has been shown and described whichsimultaneously provides high resolution, high sensitivity, and low x-raydosage to the object under investigation. The system also permits theplane of optimum focus to be varied between the target 50 and themulti-detector array 110, and provides an effective working depth offield.

Sub-sampling Technique

The following relates to a particular preferred embodiment of thepresent invention which uses the technique of sub-sampling which reducesthe computer processing overhead, and power consumption of thescanning-beam x-ray system.

Standard video quality images typically use 640×480 pixels and areupdated at 30 Hz. This usually requires a pixel sample rate of about 12Mhz. Positioning the high voltage electron beam of the x-ray tubeaccurately behind 250,000 different apertures at that rate typicallyrequires high precision and relatively high power consumption.Digitization of signals from a large array of x-ray detectors at a 12Mhz rate is similarly expensive and power intensive. Thus reduction ofthe pixel sample rate below 12 Mhz without significant reduction of thespatial or time resolution of a scanning-beam imaging system is usefulin reducing initial unit costs, operating costs due to electronic powerconsumption, and cooling requirements for the waste heat developed bythe x-ray tube.

Accordingly, a method for reducing the pixel sample rate while providingvirtually the same spatial and time resolution has been developed. Thismethod is referred to herein as "sub-sampling" and is best implementedwith the embodiment of the scanning-beam imaging system describedherein, although it could be adapted to be used with otherconfigurations. Advantages of this embodiment include reduced powerconsumption and simpler circuitry for electron beam deflection withinthe x-ray tube 10, reduced cost of fabrication of the collimation grid90, reduced complexity of the calculations needed to resolve an image ofthe object 80 and other advantages as would be obvious to those of skillin the art.

Pursuant to this embodiment a collimation grid 90 is fabricated having anumber of apertures less than the number of desired image pixels. Inother words, in sub-sampling the ratio of the number of apertures ("AP")to the number of Image Pixels ("IP") is less than one (Total AP/TotalIP<1). Preferably, AP_(x) =AP_(y) =166 rather than 500, although othernumbers are within the scope of the invention. The advantage of thisreduction from a computational point of view will become apparent below.From a manufacturing point of view, it is a much simpler structure withapproximately one-ninth the number of apertures which need to befabricated. Because this sub-sampled system requires fewer aperturesthan a fully sampled system, it is easier to fabricate grids with higherdeflection angles (i.e., the angle that the aperture makes with respectto the front face 260 of the collimation grid) without running intoproblems of having apertures 90 interfere with adjacent apertures. Thisis particularly useful when stereo grids are to be manufactured, sinceadjacent apertures in a stereo grid are directed to different detectorarrays and hence may require more physical separation than non-stereogrids to avoid aperture interference.

In the presently preferred embodiment the apertures 140 of thecollimator grid 90 preferably have a circular arrangement of maximumdimension AP_(x) rows by AP_(y) columns. For computational purposes inthe presently preferred embodiment this arrangement is treated as arectangle of dimension AP_(x) rows by AP_(y) columns with the aperturesoutside of the circular boundary contributing no information, i.e.,never being used to pass x-ray pencil beams 100.

The input face of the detector elements 160 of the multi-detector array110 are also preferably arranged in a circular array of maximumdimension DET_(x) rows by DET_(y) columns as shown in FIG. 15. Forcomputational purposes, in the presently preferred embodiment thisarrangement is treated as a rectangle of detector elements 160 ofdimension DET_(x) rows by DET_(y) columns with the detectors outside ofthe circular boundary contributing no information, i.e., always being"dark", or non-illuminated by x-rays.

The pixel sample rate is reduced by illuminating less than all of theapertures 140 of the collimation grid, if the total number of aperturesis equal to the number of image pixels, i.e., by sub-sampling.Preferably a collimation grid without the not-to-be-illuminatedapertures is used, e.g. the collimation grid includes the number ofapertures corresponding to the desired aperture to image pixel ratio. Ina collimator grid 90 having more apertures than is necessary to achievethe desired aperture to image pixel ratio, an image is formed using themulti-detector array 110, by illuminating only the collimator holes ineach row and only the collimator holes in each column that needs to beilluminated to achieve the desired aperture to image pixel ratio. Thusthe image may be built up out of image tiles of pixels corresponding tothe number of detector elements in a row in the multi-detector array 110that provides information for a single image pixel (VDET_(x)) and thenumber of detector elements in a column of the multi-detector array 100that provides information for that same image pixel (VDET_(y)) as theelectron beam 40 is scanned across the target 50. This corresponds to asub-sampling ratio of (DET_(x) ×DET_(y) /VDET_(x) ×VDET_(y)):1 which,for no sub-sampling, reduces to a sub-sampling ratio of 1:1. Thesub-sampling ratio may thus be adjusted by changing the number ofvirtual detector elements from DET_(x) to VDET_(x) in the X-direction(rows) and from DET_(y) to VDET_(y) in the Y-direction (columns). Inaccordance with a preferred embodiment, VDET_(x) =VDET_(y) =4 as shownin FIG. 36, yielding an image to aperture pixel ratio of (12×12/4×4):1,i.e., 9:1.

Where a 12×12 detector is used and the sub-sampling ratio is 144:1, theimage is fabricated from a plurality of non-overlapping images which arein effect "pasted" together--much like a photomosaic. Because real worldscintillators and detectors are usually not all perfectly andidentically responsive, the x-ray pencil beam 100 is usually notperfectly uniform, the collimation grid apertures are usually not allexactly identical with identical areas, and because a circular, ratherthan a square detector is used in the preferred embodiment, some degreeof overlap is highly desirable in order to permit averaging out detectornonlinearities and noise.

If the sub-sampling ratio is less than the detector size in image pixels(that is, less than 144:1 in this embodiment), the image will be builtup from overlapping "tiles", which are summed or averaged. If thesub-sampling ratios are not "even" multiples of the detector size (inimage pixels) or if the multi-detector array 110 is not rectangular asin the preferred embodiment there will be different numbers of samplesadded to each image pixel. To obtain a more uniform picture the valuesfrom each of the virtual detectors is normalized using differentdivisors to average the resultant values to generate each image pixel.

In the calculations that follow, VDET_(x) represents the sub-samplingvalue in the X direction (rows), and VDET_(y) represents thesub-sampling value in the Y direction (columns). For example, ifVDET_(x) =VDET_(y) =12, there is no sub-sampling. Similarly, if VDET_(x)=VDET_(y) =1, in this embodiment, the pixels are tiled. If every thirdaperture of the collimator grid 90 which has an array of apertures500×500 is illuminated, then 166×166 apertures will be scanned, i.e.,one-third in X and one-third in Y, reducing the data obtained by afactor of 9 with the 12×12 detector this will provide a sub-samplingratio of (12×12/4×4):1 or 9:1. Note that if one is only going to use166×166 apertures all of the time, there is no need for 500×500apertures and only 166×166 need be included in the collimation grid.

In accordance with one embodiment, only 1/9 of the apertures in thecollimation grid (500×500 apertures) are used or illuminated by theelectron beam 40 to generate an image. If the frame rate is keptconstant, e.g., 30 Hz, then the number of electron beam steps is reducedby 9, as is the frequency response of the circuit that drives theelectron beam. The number of scan lines is reduced by 1/3, so that theaverage horizontal beam velocity across the target is reduced by 1/3.The image reconstruction pixel rate is the same as the collimation gridaperture rate (rate at which apertures are scanned or illuminated), andis also reduced by 1/9.

In accordance with this embodiment in which the collimator grid 90includes a 500×500 array of apertures, and the multi-detector array 110includes a 12×12 array of detector elements 160, arranged such that wheneach aperture is illuminated an image pixel to aperture ratio of 1:1 isachieved, the number of detector element outputs averaged into eachimage pixel is VDET_(x) *VDET_(y). When using the sub-sampling ratio of144:1 where VDET_(x) and VDET_(y) =1, only one digitizer sample is usedfor each image pixel. The normalizing of the detector element outputssmooths out non-uniformities in the beam, the scintillators, thedetectors, and the amplifiers. The sub-sampling ratio should be set toan appropriate level for the conditions presented in order to assureacceptable image quality. This may be adjusted in accordance with theuser's preference for image quality and the conditions presented by aparticular set of circumstances.

M,N Image Reconstruction

An alternatively preferred image reconstruction method can be employedto reconstruct images along multiple focal planes. This preferred imagereconstruction method is referred to as m,n image reconstruction. FIG.58 shows a cross-sectional view through a two-dimensional array ofregularly-spaced x-ray sources and a two-dimensional array ofregularly-spaced detectors. It will be noted that there are numerousplanes parallel to the source plane 271 and detector plane 270 wheremultiple beams pass through regularly-spaced points in the plane. Theseplanes are referred to as focal planes or image planes. Theregularly-spaced points are referred to as image pixels. Each focal orimage plane comprise characteristics which differ from other focalplanes, including distance from the source, spacing of image pixels, andsize of the image plane. In accordance with the present invention, amethod is provided to reconstruct any of these image planes.

To illustrate an embodiment of this method, an array of sources,preferably a rectangular array of SOURCE_(x) by SOURCE_(y) sources on apitch λ_(s) in both the x- and y-directions, is employed with an arrayof detectors, preferably a rectangular array of DET_(x) by DET_(y)detectors on a pitch λ_(d) in both the x- and y-directions. Each sourceproduces a pencil beam of x-rays 100 which illuminates all the detectorsin the array. Each x-ray pencil beam 100 is therefore segmented into anarray of x-ray microbeams with one x-ray microbeam per detector. In thisexample, there are DET_(x),*DET_(y) microbeams per pencil beam andSOURCE_(x) *SOURCE_(y) pencil beams for a total of DET_(x) *DET_(y)*SOURCE_(x) *SOURCE_(y) microbeams.

INTENSITY(i,j,k,l) is the intensity of the x-ray illumination detectedat detector DET(i,j) when source SOURCE(k,l) is illuminated. For thisexample, i= 1,DET_(x) !, j= 1, DET_(y) !, k= 1,SOURCE_(x) !, and 1=1,SOURCE_(y) !.

Each focal plane can be described by a pair of natural numbers (integers≧1) m and n, where m*λ_(d) and n*λ_(s) are the baseline lengths of thesimilar triangles, shown in FIG. 59, which determine the location of thefocal plane. For this example, Z_(d) equals the distance from source todetector while Z_(fp) equals the distance from source to focal plane.Therefore, the distance Z_(fp) from the source plane 271 to a particularfocal or image plane which is described by the values of m,n can beexpressed as: ##EQU3##

According to this method, reconstruction of an image at a particularfocal plane m,n can be performed by creating the two-dimensional arrayIMAGE_(m),n by summing each value of INTENSITY(i,j,k,l) into image pixelIMAGE_(m),n (i*n+k*m, j*n+l*m) respectively.

The maximum x- and y-indices of array IMAGE_(m),n can be expressed as:DET_(x) *n+SOURCE_(x) *m and DET_(y) *n+SOURCE_(y) *m.

For the particular parameters of this embodiment, not all values of thenatural numbers m and n are appropriate. For example, if m and n have acommon factor (e.g. m=6, n=2) then the array IMAGE_(m),n will besparsely filled. The image reconstructed using m=6, n=2 is equivalent tothe image reconstructed using m=3, n=1. Although the array IMAGES₆,2will have four times as many elements as IMAGE₃,1 only one-fourth of theelements in IMAGE₆,2 will be non-zero. Removing the all-zero rows andcolumns in IMAGE₆, 2 yields IMAGE₃,1.

Referring to FIG. 59, in this example it will also be noted that, e.g.,doubling or tripling both baselines of the similar triangles does notchange the location of the resulting focal plane.

The pitch of the image pixels at the focal plane λ_(fp) can be expressedas follows: ##EQU4##

Referring to FIG. 59, it will be noted that every mth detector in the x-and y-directions is used to reconstruct any particular image pixel.Therefore, there are, on average, DET_(x) *DET_(y) /m² microbeams perimage pixel. Since the total number of microbeams in this example isDET_(x) *DET_(y) *SOURCE_(y) *SOURCE_(y), the number of image pixels canbe expressed as: ##EQU5##

Due to partial image reconstruction around the perimeter of the image,the number of fully reconstructed image pixels is slightly lower thanthe above number and the total number of fully and partiallyreconstructed image pixels is slightly higher than the above number.

In this example, when the size of the source array is SOURCE_(x) *λ_(s)by SOURCE_(y) *λ_(s), the size of the field of view (at the focal plane)can be expressed as: ##EQU6##

The m,n image reconstruction method is more flexible than the previouslydescribed reconstruction methods. As FIG. 58 shows, m,n imagereconstruction can generate a wide variety of focal planes at numerouspositions between the source and detector planes. Many of the focalplanes have a small pitch between image pixels which can be used toproduce images with high spatial resolution.

The ability to reconstruct a wide variety of focal planes can be used tomove the focal plane with respect to the source and detector by simplyselecting a suitable image plane near the region of interest of theobject to be imaged.

The m,n image reconstruction method can also be used to increase theeffective depth of field of an image by simultaneously reconstructingmultiple focal planes around a region of interest. The reconstructedplanes can be combined to produce a single image with high spatialresolution over a larger range of distances from the x-ray source plane.The multiple reconstructed planes can be combined, for example, byadding together only the high spatial frequency components from eachreconstructed plane.

System Description

FIGS. 18-25 are functional block diagrams of a preferred stereoscopicscanning-beam x-ray imaging system according to the present invention.FIG. 17 depicts the manner in which FIGS. 18-25 can be arranged tocreate a single block diagram of this presently preferred system. Formedical applications the x-ray source and the multi-detector arrays arepreferably mounted on a movable C-arm with the x-ray source mountedabove an adjustable patient table and the multi-detector arrays locatedbelow the table.

FIG. 18 includes a representative block diagram of high-voltage terminal803, which is part of x-ray source 798. High-voltage terminal 803, whichis preferably contained in a housing (not shown) includes electricalcomponents for producing and controlling the electron beam 1240. Thehigh voltage necessary to power the x-ray source is supplied to the highvoltage terminal from an external adjustable high-voltage power supply790 by a single cable 1010. All of the electronic components in the highvoltage terminal 803 preferably float at the output potential of thehigh-voltage power supply. The unique construction of the high voltageterminal results in only this single electrical connection to the highvoltage terminal. All other data and power transmission to and from thehigh voltage terminal is preferably accomplished via fiber optic linksand via an isolation transformer. A more detailed description of thestructure and operation of the presently preferred high voltage terminalis provided in copending U.S. patent application Ser. No. 08/386,884,which has been incorporated herein by reference in its entirety.

Because of the high operating voltages with respect to ground in thehigh-voltage terminal 803 (-100 kV to -120 kV), the housing ispreferably filled with an electrically insulating medium, preferablypressurized SF₆ (sulphur hexafluoride) gas, to electrically isolate thehigh voltage terminal 803 from its surroundings.

High voltage isolation transformer 744 supplies power for the circuitryin the high-voltage terminal 803. The secondary winding 1271 ofisolation transformer 744 is part of the high-voltage terminal 803,while the primary winding 1270 is separated from the terminal by a gapfilled with the pressurized SF₆ gas. The primary winding 1270 preferablyforms a part of the high voltage terminal housing and is supplied withpower from the high voltage terminal power supply driver 772 located inthe C-arm cart. The preferred construction of the isolation transformeris described more fully in copending U.S. patent application Ser. No.08/386,884, which has been incorporated herein by reference in itsentirety.

In addition to the components necessary for generating and controllingan electron beam, the high voltage terminal preferably includescomponents for monitoring certain parameters located within the highvoltage terminal. The monitored information is preferably communicatedoutside the housing via fiber optic cables. The circuitry for convertingthe electrical signals to light signals and the light signals toelectrical signals is described more fully in connection with thedetailed description of FIGS. 40A-E and 41A-B and is enclosed within thehigh voltage terminal housing.

A pressure sensor 722 preferably monitors the pressure of the SF₆ gas inthe housing to ensure adequate electrical isolation. Additionally, thetemperature of the pressurized SF₆ gas is preferably monitored bytemperature sensor 724. This information is transmitted via multiplexer732 and I/O controller 762 to control computer 890. If the pressuredrops below a predetermined threshold or the temperature increases abovea predetermined threshold, the control computer will shut the systemdown.

The electric current from the high voltage power supply is preferablysensed by passing the current through a beam current sensor 736 whichprovides information to a current sense monitor 788 which is preferablylocated in the C-arm cart.

In addition the heater controller 734, which controls the heater 726located in the electron gun 728, provides information concerning heatercurrent and voltage to multiplexer 732 for transmission to the controlcomputer 890.

The voltage of the electron grid 730, located in front of the emittingface of electron gun 728, is controlled by grid controller 738. Thevoltage level of electron grid 730 can preferably be varied between zeroand -2000 V with respect to the cathode to adjust the current of theelectron beam 1240, thereby controlling the x-ray flux emitted by target1250. When the electron grid is at a potential of approximately -2000 V,the electron beam is effectively shut off. The beam-on control signal is1024 instructs the grid controller 738 to apply either -2000 V to thegrid to turn the electron beam off, or to apply the beam-on grid voltageset via fiber optic link 1030 to the grid to turn the electron beam onto a preset current. The grid controller 738 also relays the beam-on andbeam-off grid voltages to the multiplexer 732. Fault conditions in thegrid controller will trip the high voltage terminal fault latch 742which will shut the electron beam off by turning off heater 726 viaheater controller 734 and setting the voltage of electron grid assembly730 to -2000 V via grid controller 738. The entire x-ray source willalso be shut down via fiber-optic cable 1020 and fail safe controller760.

The status information from various components within high voltageterminal 803, which are input to multiplexer 732, are transmitted to I/Ocontroller 762. Multiplexer 732 includes a voltage to frequencyconverter which drives an LED for conversion of the electrical statussignals from a selected component into light pulses and transmits thesesignals to I/O controller 762 via fiber-optic cable 1016. I/O controller762 controls the sequence of transmission of each component's statusinformation sent via multiplexer 732 by sending a channel select signalto multiplexer 732 via fiber-optic cable 1018.

Ion pump 720 maintains the vacuum within the x-ray source 798. Ion pump720 is powered by ion pump power supply 758, which in the preferredembodiment is located within C-arm cart 811. Ion pump power supply 758also has an output which indicates the vacuum pressure to controlcomputer 890 via I/O controller 762.

Referring now to FIG. 19, electron gun 728 emits a beam of electrons1240 toward grounded target 1250 which preferably passes through focuscoils 746 and deflection coils 748 to focus and position electron beam1240 at a desired location on target 1250. The deflection coils 748 aimelectron beam 1240 at a specific location on the surface of target 1250.Target 1250 emits x-rays 1241 at the spot illuminated by electron beam1240. Infra-red temperature sensor 750 constantly monitors thetemperature of target 1250 for abnormal increases in temperature causedby malfunctioning of the beam scanning. If infra-red temperature sensor750 detects an over-temperature condition, it trips the fail-safecontroller 760 to shut down the x-ray source. To verify proper operationof the temperature sensor, an infra-red test source is provided whichcan be activated by the control computer to simulate an over-temperaturecondition to verify that the infra-red temperature sensor will detect afault and shutdown the x-ray source.

A cooling chamber 754 for cooling the target 1250 is preferably locatedbetween the target 1250 and the collimator 1290. The coolant fromcooling chamber 754 is preferably circulated through a heat exchanger756 (FIG. 24) preferably housed within the C-arm. Since the collimator1290 may come in contact with the patient during imaging procedures, thecollimator 1290 is preferably monitored for excessive temperatures bycollimator temperature sensor 752. In this preferred embodiment,collimator temperature sensor 752 checks for temperatures in excess of40° C. If the temperature exceeds this threshold, the fault iscommunicated to the fail-safe controller 760, which shuts down the x-raysource. To verify proper operation of the temperature sensor, a testheater is provided which can be activated by the control computer tosimulate an over-temperature condition to verify that the temperaturesensor will detect a fault and shutdown the x-ray source.

FIG. 23 is a block diagram comprising beam controller 796 and a portionof the C-arm cart. The beam controller interface 794 receives data fromthe tube controller via a high speed fiber optic link 1000.Consequently, beam controller interface 794 includes the light signal toelectrical signal conversion circuitry described more fully inconjunction with FIGS. 40A-E and 41A-B.

Beam controller 796 preferably controls the focus coils 746 through twoseparate drivers, a static focus driver 774 and a dynamic focus driver776. Static focus driver 774 is preferably set only once for a givenoperating voltage of the high voltage power supply. The dynamic focusdriver 776 adjusts the precise focussing of the electron beam 1240 as itscans across target 1250.

Beam controller 796 preferably controls the deflection coils 748 throughfive separate drivers: x-deflection driver 778, x-step driver 780,y-deflection driver 782, 45° stigmator driver 784, and 0° stigmatordriver 786.

The x-deflection driver 778 communicates a conventional linear inputpattern to the deflection coils via wires 1046 to drive the electronbeam horizontally across the target whereas the x-step driver 780communicates a novel sawtooth input signal to the deflection coils 748via wires 1048. The net effect is a stepped movement of the electronbeam across the target. The y-deflection driver 782 communicates aconventional y-deflection pattern to the deflection coils 748 via wires1050 to drive the electron beam 1240 vertically across the face of theanode. The 45° stigmator driver 784 and the 0° stigmator driver 786 andtheir respective coils correct for aberrations in the electron beam spotto maintain a circular spot on the target. More detailed informationabout these circuits can be found in copending U.S. patent applicationSer. No. 08/386,884, which has been incorporated herein by reference.

Current sense monitor 788 is preferably used to monitor the output ofany of the beam controller drivers to verify their correct operation aswell as to measure the electron beam current as previously discussed.

A failure in the deflection system could result in the electron beam notscanning across the target in the x direction or the y direction. Thiscould result in thermal damage to the target. Deflection fault sensor770 preferably receives x-scan and y-scan monitoring information fromx-deflection driver 778 and y-deflection driver 782. Deflection faultsensor 770 preferably transmits a fault status signal to fail-safecontroller 760 via fiber-optic cable 1072. If a deflection faultcondition occurs, fail-safe controller 760 will shutdown the x-raysource.

FIGS. 23 and 24 include a functional block diagram of the componentspreferably housed in the C-arm cart 811. Power is preferably supplied tothe C-arm cart 811 from a 208 volt 3-phase AC power supply via cable7063. DC power is fed from the c-arm cart to the beam controller 796 viacable 1078.

I/O controller 762 (FIGS. 23 and 24) preferably communicates with thecontrol computer 890 via high speed fiber-optic cables 1002 and 1004 andincludes the electrical to light and light signal to electrical signalconversion circuitry described more fully in conjunction with FIGS.40A-E and 41A-B. Beam controller interface information including gridvoltage, static focus current, current sense select, current sensesample select information and current sense sample information, istransmitted to beam controller interface 794 from I/O controller 762 viacable 1080.

As discussed, fail-safe controller 760 preferably receives and monitorsstatus information from various components of the system and is designedto disable the system upon detection of a potential safety problem. Ifthe fail-safe controller 760 detects such a potential problem, it willpreferably: (1) signal the grid controller 738 to disable (turn off) theelectron beam; (2) shut down the high-voltage power supply 790; and (3)shut down the static focus driver 774 to defocus the electron beam.

In the preferred embodiment, fail-safe controller 760 receives faultstatus signals from: heat exchanger 756 via wire 1120; collimatortemperature sensor 752 via wire 1066; IR target temperature sensor viawire 1068; high-voltage terminal fault latch 742 via fiber-optic cable1020; and deflection fault sensor 770 via fiber-optic cable 1072.Fail-safe controller 760 also relays the fault status signals to thecontrol computer via I/O controller 762 so that fault conditions may bedisplayed and logged by the control computer.

High-voltage power supply 790 is preferably located on C-arm cart 811.The signal to turn on the high-voltage power supply 790 is sent from theI/O controller 762 to the high voltage supply 790 via wire 1144. Voltagesetpoint is sent to high-voltage power supply 790 via wire 1140, andcurrent limit is sent via wire 1142. Voltage monitoring signal is sentto the I/O controller from the high voltage supply via wire 1146 andcurrent monitoring signal is sent via wire 1148.

In FIG. 20, maneuverable and locatable catheters 1285 are shown insertedinto patient 1280. The proximal end 1284 of catheters 1285 arepreferably connected to catheter connector 970. Catheter connection 970is preferably connected to a multi-channel photomultiplier tube 900(FIG. 22) through fiber-optic cable 980.

FIG. 20 also functionally diagrams preferred right 822 and left 1522detectors of the present invention. Since both detectors in FIG. 20function in a similar fashion, only the right detector 822 will bediscussed in detail. The components bearing a number having the samelast two digits perform the same function.

Scintillator array 802 preferably comprises ninety-six elements and inresponse to x-ray photons generates visible light energy which istransmitted to photomultiplier tube 806 comprising ninety-six channelsvia a tapered fiber-optic bundle 804. The photomultiplier tube 806converts the received light energy into electrical signals which aresent to signal conditioner 810 via 96 separate electrical connections836. These signals are referred to herein as raw partial image pixelinformation. The multi-detector array preferably comprises ascintillator array 802, fiber-optic taper 804 and photomultiplier tube806. It should be noted that while the preferred embodiment includes 96channels, more or less than that number are within the spirit and scopeof the present invention. Photomultiplier tube 806 is powered byphotomultiplier tube power supply 808.

Signal conditioner 810 is preferably comprised of 48 circuit boards1343. Each circuit board 1343 comprises two sets of signal conditioningamplifier circuits 1830, with each signal conditioning amplifier circuit1830 feeding its output to a corresponding discriminator 1832. Thus 96sets of signal conditioning amplifier circuits 1830 and discriminators1832 are employed, with each set paired to a correspondingphotomultiplier tube channel. The signal conditioner 810 outputsninety-six separate signals for every step of the electron beam. Thisinformation is referred to as the partial image pixel information.

The outputs of the signal conditioners are preferably input into thebeam alignment extractor 816. Beam alignment extractor 816 processes theinformation from each position of the electron beam on the target andsends processed alignment data to data transmitter 818. Clock signalsare sent to the beam alignment extractor from data receiver 812.

Beam alignment extractor 816 sends the partial image pixel informationfrom signal conditioner 810 to image reconstruction engine 814. Fordiagnostic purposes, the partial image pixel information sent fromsignal conditioner 810 may be modified by the beam alignment extractor816 before it is sent to the image reconstruction engine 814. Imagereconstruction engine 814 processes the partial image pixel informationand sends image pixel data to data transmitter 818. The imagereconstruction engine 814 receives clock signals from data receiver 812via electrical connection 834.

The detector controller 805 (FIG. 21) for the detectors 822 and 1522preferably transmits and receives optical signals to and from thedetectors. Right receiver 880 receives image pixel data and beamalignment data from the right detector 822 through high-speedfiber-optic cable 826. Right detector 822 transmits this data through adata transmitter 818 (FIG. 20), which preferably includes circuitry forconversion of the signals from image reconstruction engine 814 and beamalignment extractor 816 into a serial signal. This serial signal isconverted into light pulses using an LED. Right receiver 880 alsocomprises a light detector and related circuitry for receiving anddecoding the light pulse from a serial signal into parallel signals. Thebeam alignment data is transmitted to control computer 890. The imagepixel data is preferably transmitted to frame buffer 872. The leftreceiver 846 operates in a similar fashion to receive image pixel dataand beam alignment data from the left detector 1522.

Right transmitter 886 comprises circuitry for converting parallelsignals into serial signals. Right transmitter 886 receives, among othersignals, signals to set channel gains and threshold levels from controlcomputer 890. Right transmitter 886 also receives clock signals frombeam deflection lookup table 918 (FIG. 25). These signals are convertedinto serial signals which are then transmitted as light pulses to theright detector 822 through high-speed fiber-optic cable 824. Right datareceiver 812, which contains a light detector and circuitry to convertlight pulses into parallel signals receives these signals. The signal toset channel gain is transmitted to the signal conditioner 810 throughwire 828. The left transmitter 848 operates in a similar manner tocommunicate control signals to the left detector 1522.

Image pixel data transmitted to right frame buffer 872 is subsequentlytransmitted to video processor 858 where in a stereoscopic system, it ispreferably combined with image pixel data from left frame buffer 850.Brightness and contrast information are transmitted from the right framebuffer 872 and from the left frame buffer 850 to control computer 890.This information is used to set the output of the x-ray source foroptimal image quality and x-ray exposure control. Control computer 890transmits information to the video processor 858 for annotation of theimage display. The output of video processor 858 is preferably sent toimage display monitor 862 where the image is displayed.

Control computer 890 preferably controls the operation of the system viadetector controller 805, tube controller 807, and beam controller 796.Control computer 890 may receive operator instructions from inputsources such as keyboard 894, trackball 896, and control panel 898. Theoperator receives system information from the control computer throughcontrol monitor 892 and speaker 899.

Referring to FIG. 22, catheter processor 809 preferably receivesinformation from up to eight catheters 1285 via fiber-optic cables 980.The light pulses received through fiber-optic cables 980 are preferablydetected by the catheter multi-channel photomultiplier tube 900. Thecatheter multi-channel photomultiplier tube 900 is powered by powersupply 906. The information received by the catheter multi-channelphotomultiplier tube 900 is preferably sent to catheter signalconditioning circuit 902 via electrical connection 910. The cathetersignal conditioning circuit 902 outputs data to the catheter dataextractor 904 via electrical connection 908. The catheter informationfrom catheter data extractor 904 is transmitted to the control computer890.

Tube controller 807 transmits data to and from the I/O controller 762and the beam controller 796 to control the operation of x-ray source798. Tube controller 807 preferably comprises beam deflection lookuptable 918, programmable scan controller 920, beam transmitter 916, I/Otransceiver 964, and I/O fault latch 958.

Programmable scan controller 920 is preferably set by control computer890 to produce a particular scan. These setting may include, forexample, scan rate, serpentine or raster scan, and round or square scan.Programmable scan controller 920 transmits a sequence of desired beampositions to beam deflection lookup table 918. For each desired locationof the electron beam, the beam deflection lookup table preferablycontains values for deflection and focus necessary to produce a wellfocused spot at the correct location on the target. The data in the beamdeflection lookup table 918 is preferably programmed by control computer890.

Data from the beam deflection lookup table 918 is preferably sent tobeam controller interface 794 via beam transmitter 916 and high-speedfiber-optic link 1000. This data includes: (1) current sense samplesignals; (2) dynamic focus; (3) x-step; (4) x-deflection; (5)y-deflection; (6) 45° is stigmator; (7) 0° stigmator; and (8) "beam onrequest" signals. Preferably, approximately every 1.28 microseconds, anew set of data is sent from the beam deflection lookup table 918 to thebeam controller interface 794.

I/O transceiver 964 provides the communications link between the controlcomputer 890 and the I/O controller 762. Control computer 890 sends dataand control signals to I/O controller 762. Information from the x-raysource 798 is sent to the control computer 890 via I/O transceiver 964.

If a fault condition occurs during the transmission of information frombeam transmitter 916 to beam controller interface 794, deflection faultsensor will detect the fault and shut the x-ray source down via failsafe controller 760. If a fault condition occurs during the transmissionof information from I/O transceiver 964 to I/O controller 762, the I/Ocontroller will detect the fault and shut the x-ray source down via failsafe controller 760. If a fault condition occurs during the transmissionof information from I/O controller 762 to I/O transceiver 964, I/Otransceiver 964 will set the I/O fault latch 958 which will disablecommunications via fiber-optic cables 1000 and 1002. This will bedetected as faults by the deflection fault sensor and by the I/Ocontroller which will shut the x-ray source down as described above.

Real-Time Eye

FIG. 26 depicts "real-time eye" assembly 402 which comprises themulti-detector array according to a presently preferred embodiment ofthe present invention. X-rays enter through x-ray window 404 in leadshield 406. X-ray window 404 is preferably circular and about 1.91 cm(0.75 in) in diameter to permit x-rays coming from the apertures 140 ofthe collimation grid 90 to strike the multi-detector array 110 whileattenuating scattered x-rays. A light shield 408 is preferably providedto shield the eye from ambient light. The light shield 408 may be madeof a thin sheet of aluminum or beryllium chosen to attenuate lightwithout substantially attenuating the x-rays, and is preferably 0.0125cm thick. The multi-detector array assembly 402 is preferably enclosedin a light-tight outer detector housing 418 to minimize stray light fromgenerating noise. Three centering screws 422 are provided for planar andlinear alignment. Rotational alignment in one embodiment is achieved byrotating outer detector housing 418 with respect to PMT mount 426.

Scintillator array 112 is preferably mounted beneath the x-ray window404. Scintillator array 112 is preferably comprised of 96 scintillatorelements 170 arranged in a pseudo-circle, with each scintillator element170 preferably cut to a square horizontal cross-section. The length ofthe individual scintillator elements 170 are preferably about 0.50 cmand the front input faces are preferably 0.135 cm×0.135 cm. Thescintillator elements 170 are preferably YSO, LSO or BGO but otherscintillating materials may also be used.

For a suitably reduced decay time for its light output in thisapplication (to about 50 nsec), BGO needs to be heated to approximately100° C. When using BGO, the scintillator array is located near heatingelement 410 for use with a BGO scintillator. If a BGO scintillator isused heating element 410 may be a resistive heating element designed tokeep the BGO scintillator crystal array 112 at an operating temperatureof about 100° C. Accordingly a resistive heating element may beprovided, as shown in FIG. 26. YSO is preferably used as thescintillator material, thereby avoiding the need for a heater.

A fiber-optic imaging taper 412 of the preferred multi-detector array110 directs light photons emerging out of the bottom 414 of thescintillator crystal array 112 to a 96 channel photomultiplier tube(PMT) 416. A presently preferred fiber-optic imaging taper 412 isavailable from Collimated Holes of Campbell, Calif. and has a circularinput aperture of diameter 2.03 cm (0.8 in) and a circular outputaperture of diameter 3.38 cm (1.33 in). Taper 412 matches eachscintillator crystal pitch dimension (0.06") to that of the PMT416(0.10"), i.e., it has a magnification of 1.667 times. High viscosityoptical coupling fluid available from Dow Corning (Type 200) with arefractive index approximately matching that of the glass may be used atthe two faces of the taper as an optical coupling medium to maximize thelight transfer efficiency from the scintillator crystals 170 to thetaper 412 and from the taper 412 to the PMT input face 424.

Photomultiplier tube 416 is preferably a 96 channel tube (one channelcorresponding to each scintillator crystal 170) model number XP 1724Aavailable from the Philips Corporation. Photomultiplier tube 416preferably has a fiber-optic face plate so that the spatial arrangementof the scintillator crystal array 112 is accurately carried through tothe PMT photocathode located in the PMT on the other face of thefaceplate. An x-ray photon striking one of the scintillators 170produces many light photons which are coupled to the PMT photocathode.This produces a corresponding electron pulse at the photocathode and thepulse is amplified in one channel of the PMT dynode structure up to1,000,000 times.

Referring to FIGS. 27 and 28, the front face of PMT 416 includes a glasswindow 1340 which extends beyond the PMT encapsulation by 0.1 mm. 96photo-cathode elements 1339 are arranged in a pseudo-circular array inthe center of the front face of PMT 416. Each photo-cathode element issquare in shape with dimensions of 2.54 mm×2.54 mm. PMT 416 is attachedto the PMT mount by means of attachment bolts set into PMT 416 at boltholes 1374.

This pseudo-circular array of 96 photo-cathode elements creates alight-sensitive circular area 1338 on the PMT 416 with a diameter of30.5 mm. It is this light sensitive area 1338 that interfaces with thetapered fiber-optic bundle 412. Each PMT photocathode element 1339 has acorresponding electrical output connector 1342. When light photons reachthe PMT 416, the photocathode elements 1339 generates raw partial imagepixel signals which is output at PMT connector 1342. The raw partialimage pixel signals are transmitted via PMT connector 1342 to signalconditioner 810. Further details of a presently preferred real time eyeassembly can be found in co-pending U.S. patent application Ser. No.08/387,292 which has been incorporated herein by reference in itsentirety.

Signal Conditioner

Signal conditioner 810 preferably converts the 96 outputs of PMT 416into 96 pulse trains with each pulse in the pulse train corresponding toa single x-ray photon arriving at the corresponding scintillator element170. Signal conditioner 810 is preferably comprised of 48 circuit boards1343. Each circuit board 1343 comprises two sets of signal conditioningamplifier circuits 1830, with each signal conditioning amplifier circuit1830 feeding its output to a corresponding discriminator 1832. Thus 96sets of signal conditioning amplifier circuits 1830 and discriminators1832 are employed, with each set paired to a correspondingphotomultiplier tube channel. The signal conditioner 810 outputsninety-six separate pulse trains for every step of the electron beam.This information is referred to as the partial image pixel information.

The signal conditioning amplifiers 1344 shape and amplify the rawpartial image pixel signals from the photomultiplier tube, and output apulse train of partial image pixel signals to the beam alignment andimage reconstruction boards. To even out any performance variationsbetween the individual photomultiplier tube channels, a separate gainsignal is sent to each of the signal conditioning amplifier circuits1830. However, the same threshold signal is sent to each discriminator1832.

FIGS. 29A-D are a circuit diagram of a preferred signal conditioningamplifier circuit 1830. Raw partial image pixel signals from a singlephotomultiplier tube channel are input to the signal conditioningamplifier circuit 1830 via input line 1834. Signal conditioningamplifier circuit 1830 is preferably AC coupled to eliminate offsetdrift problems. The AC coupling low frequency cut-off is high, e.g., 30Mhz, so that the pulse is differentiated. This eliminates the need for aDC restorer circuit to keep the baseline reference voltage constant asthe pulse rate varies. Clamping diodes 1848 provide voltage protectionfor the amplifiers within the signal conditioning amplifier circuit1830.

Signal conditioning amplifier circuit 1830 preferably comprises threestages of current amplification. The input partial image pixel signalsare coupled through a coupling capacitor 1842 to a fixed gain firststage amplifier 1836. The output of the first stage amplifier is fed toa variable gain second stage amplifier 1840, which receives gain controlsignals applied over input line 1844. The output from the variable gainsecond stage amplifier 1840 is fed to a fixed gain third stage amplifier1838, which sends an amplified partial image pixel waveform to thediscriminator via line 1846. Supply voltages of +5V and -5V are appliedto the each amplifier stage within the signal conditioning amplifiercircuit 1830. Each of the 96 signal conditioning amplifier circuits 1830function similarly to process raw partial image pixel signals from itscorresponding photomultiplier tube-channel.

Referring to FIG. 30, the discriminator 1832 essentially digitizes thepartial pixel information by comparing the amplified partial image pixelwaveform from a signal conditioning amplifier circuit 1830 with athreshold value and producing a high or low value depending on whetherthe threshold value is crossed. This high or low value corresponds towhether an x-ray photon was detected or not. FIG. 30 diagrams thepreferred input and output connectors for each pair of discriminators1832 located on a single circuit board 1343. The amplified partial imagepixel waveforms from two sets of signal conditioning amplifier circuits1830 are coupled to the discriminators 1832 via input lines 1846 and1847. Surface mount ferrite bead inductor 1850 are preferably employedto filter noise from the input waveforms. Digitized output pulses fromthe discriminators 1832 are output via output lines 1848 and 1849.

FIG. 31 is a circuit diagram of a preferred discriminator 1832. Theamplified partial image pixel waveforms from the signal conditioningamplifier circuit 1830 are input, via input line 1846, to a comparator1854 which provides a constant amplitude output pulse regardless of theamplitude of its input. The threshold reference signal, applied tocomparator 1854 via input line 1852, is preferably set to a value whichis slightly higher than the amplifier noise output level so that it willnot trigger on the noise level. The supply voltage inputs for thecomparator 1854 are preferably set at +5V and -5V.

The preferred comparator 1854, a standard LT1016 comparator availablefrom Linear Technology, functions as both a comparator and a register.Comparator 1854 generates a latched output which is preferably coupledto a pulse stretching circuit 1856, which is comprised of a circuitrydiode, grounded resistor and a capacitor. The pulse stretching circuit1856 allows the comparator 1854 to generate output pulses approximately29 nanoseconds wide. The output pulse from the comparator is preferablyfed to a divide-by-two counter 1858, to reduce the frequency of theoutput pulses. No information is lost since subsequent circuits countthe edges of this pulsed output. The pulsed output, containing digitizedpartial image pixel signals, are output from the divide-by-two counter1858 to the next processing stage via output line 1848. Each of the 96discriminators 1832 function in a similar manner to process partialimage pixel waveforms from its corresponding signal conditioningamplifier circuits 1830.

FIGS. 32A-G diagram the DACs (digital-analog converters) whichpreferably provide the gain and threshold control signals for the signalconditioner 810. The 12 gain control DACs 1860 each receive serialcontrol data from the control computer, and output a total of 96parallel analog gain control signals through 48 interface connectors1862 (FIGS. 33A-I) to corresponding signal conditioning amplifiercircuits 1830. A threshold control DAC 1864 receives digital controldata from the control computer, and outputs a single threshold referencesignal which is sent to all 96 discriminators 1832. Threshold controlDAC 1864 feeds its output threshold reference signal to a bufferamplifier 1866 (FIG. 34), which provides the power to drive all 96threshold reference inputs to the individual discriminators 1832.

FIGS. 35A-D diagram the connectors 1868 between the output of thediscriminators 1832 and the image reconstruction and beam alignmentcircuit boards.

Image Reconstruction with Sub-sampling

The presently preferred image reconstruction method utilizes thesub-sampling method to process the detected information. Preferably thesub-sampling method is employed in a reverse geometry scanning beamx-ray system utilizing a sub-sampling ratio of 9:1 with a multi-detectorarray 822 including ninety-six detector elements arranged in apseudo-circle.

FIG. 36 is a diagram of a 12 by 12 logical array 823 of detectorelements. The logical array includes both active detector spaces 642 andinactive detector spaces 640. In the presently preferred imagereconstruction method the 96 active detector element spaces each includea detector element and form an active logical array 822 which occupy thecenter spaces of a 12 by 12 logical array arranged in a symmetricalpattern about the horizontal midline and the vertical midline. Theremaining 48 logical detector spaces of the array are inactive detectorsand preferably do not include a detector element. In the preferredembodiment the inactive detector spaces do not output real informationabout the object.

To generate an image pixel, the processed x-ray intensity valuesdetected by the multi-detector array 110 for each x-ray micro-beampassing through that image pixel IP are summed and output to a videomonitor. For image reconstruction using a sub-sampling ratio of 1:1 eachlogical detector element of the logical array is capable of providinginformation about each image pixel in the object. For imagereconstruction with a sub-sampling ratio of x:1, where x is a numbergreater than 1, less than all of the logical detector elements arecapable of contributing information about a particular image pixel. Theactual number capable of contributing information will depend on theparticular sub-sampling ratio selected. With a presently preferredsub-sampling ratio of 9:1 in the presently preferred embodiment, only 16logical detector elements of the 144 logical detector element logicalarray 823 will provide information about any particular image pixel.

In the sub-sampling method with a sub-sampling ratio of 9:1 the logicalarray 823 includes sixteen virtual detectors, e.g., 644, 646, 648 and649. In this embodiment the virtual detectors each include 9 logicaldetectors arranged in a 3 by 3 array. Alternatively, if a sub-samplingratio of 4:1 were used, there would be 36 virtual detectors, eachincluding 4 logical detector elements. Using a sub-sampling ratio of 1:1there would be 144 virtual detectors each including 1 logical detectorelement.

Each of the 16 logical detector elements used to reconstruct a singleimage pixel using a sub-sampling ratio of 9:1 are preferably situated indifferent virtual detectors. In this embodiment, each virtual detectorcontributes partial image pixel information for nine different imagepixels. Complete image pixel information is obtained by combining theinformation from the logical detectors in the same virtual arraylocation from all 16 virtual detectors.

The presently preferred image reconstruction method utilizes a novelstring method. In the string method there is one string for each logicaldetector in a virtual detector. For example, using the preferredsub-sampling ratio of 9:1 there are nine strings. Referring to FIG. 36,each of the virtual array locations of the virtual detectors have beenassigned numbers 1 through 9. String 1 includes all of the logicaldetectors assigned the number 1. String 2 includes all of the logicaldetectors assigned the number 2. And so on. Each row of the logicalarray 823 is assigned a number, 1 through 12, going from top to bottom.Each column of the logical array is assigned an alpha character, Athrough M, going from left to right. Naturally the use of right, left,top and down is relative and the particular orientation selected ismerely to more easily explain the method of image reconstruction.

FIG. 37 is a diagram of the process flow used in the string method.Since the string method is the same for all strings, the method will bedescribed in detail with regard to only string 1.

String 1 is comprised of the following logical detectors, M1, J1, F1,C1, M4, J4, F4, C4, M7, J7, F7, C7, M10, J10, F10 and C10. Assuming theelectron beam moves from left to right for each row and top to bottom,and the sub-sampling ratio is 9:1, using the preferred multi-detectorarray 110, the first logical detector element that is capable ofreceiving information for a particular image pixel is M10 (FIG. 36).When the electron beam is positioned behind the next aperture (one holeto the right from the view of the output face of the collimator), thesecond logical detector capable of receiving information about that sameimage pixel is J10. And so on.

Referring back to FIG. 37, the partial pixel information for each imagepixel is preferably processed for each string in accordance with thefollowing method. The method will first be described in accordance withthe embodiment in which each of the logical detectors are active and asub-sampling ratio of 9:1 is selected resulting in a collimatorincluding 167 rows and 167 columns. The description begins when theelectron beam is positioned behind aperture AP₅₀,50. For purposes ofthis description image pixel IP₁ is located along the axes of the x-raymicro-beam detected by logical detector M10. Further, when the electronbeam is described as being located behind a particle aperture, it meansthat the electron beam is aimed at the intersection of target layer andthe axis of the x-ray pencil beam (which is aimed at the center of themulti-detector array) formed by that aperture.

As the electron beam is being positioned behind AP₅₀,50, subcounter M10is reset to zero. While the electron beam is positioned behind AP₅₀,50the partial image pixel information detected by logical detector M10 isinput to subcounter M10. As the electron beam is positioned behind thenext aperture in the same collimator row AP₅₀,51, the informationcontained in subcounter M10 is moved to subcounter J10. While theelectron beam is positioned behind the next selected aperture AP₅₀,51 anx-ray micro-beam will pass through IP₁ and strike logical detector J10.The partial image pixel information detected by logical detector J10will be input to and added to the contents of subcounter J10.

As the electron beam is positioned behind the next aperture in the samecollimator row AP₅₀,52, the information contained in subcounter J10 ismoved to subcounter F10. While the electron beam is positioned behindthe next selected aperture AP₅₀,52, an x-ray micro-beam will passthrough IP₁ and strike logical detector F10. The partial image pixelinformation detected by logical detector F10 will be input to and addedto the contents of subcounter F10.

As the electron beam is positioned behind the next aperture in the samecollimator row AP₅₀,52, the information contained in subcounter F10 ismoved to subcounter C10. While the electron beam is positioned behindthe next selected aperture AP₅₀,53, an x-ray micro-beam will passthrough IP₁ and strike logical detector C10. The partial image pixelinformation detected by logical detector C10 will be input to and addedto the contents of subcounter C10.

As the electron beam is positioned behind the next aperture in the samecollimator row AP₅₀,54, the information contained in subcounter C10 ismoved to a FIFO register. The reason for this is that because of thegeometry of the preferred system, when the electron beam is positionedbehind the next selected aperture AP₅₀,54, no x-ray micro-beam will passthrough IP₁ and strike any logical detector in the array until theelectron beam is moved to the next row. In accordance with thisembodiment, no x-ray micro-beam will pass through IP₁ and strike anylogical detector until the electron beam is positioned behind AP₅₁,50.

As the electron beam is being positioned behind AP₅₁,50, subcounter M7is loaded with the partial image pixel information stored in the FIFOcorresponding to the partial pixel information output from C10. Whilethe electron beam is positioned behindAP₅₁,50 the partial image pixelinformation detected by logical detector M7 is added to the contents ofsubcounter M7.

As the electron beam is positioned behind the next aperture in the samecollimator row AP₅₁,51, the information contained in subcounter M7 ismoved to subcounter J7. While the electron beam is positioned behind thenext selected aperture AP₅₁,51 an x-ray micro-beam will pass through IP₁and strike logical detector J7. The partial image pixel informationdetected by logical detector J7 will be input to and added to thecontents of subcounter J7.

As the electron beam is positioned behind the next aperture in the samecollimator row AP₅₁,52 the information contained in subcounter J7 ismoved to subcounter F7. While the electron beam is positioned behind thenext selected aperture AP₅₁,52, an x-ray micro-beam will pass throughIP₁ and strike logical detector F7. The partial image pixel informationdetected by logical detector F7 will be input to and added to thecontents of subcounter F7.

As the electron beam is positioned behind the next aperture in the samecollimator row AP₅₁,53, the information contained in subcounter F7 ismoved to subcounter C7. While the electron beam is positioned behind thenext selected aperture AP₅₁,53, an x-ray micro-beam will pass throughIP₁ and strike logical detector C7. The partial image pixel informationdetected by logical detector C7 will be input to and added to thecontents of subcounter C7.

As the electron beam is positioned behind the next aperture in the samecollimator row AP₅₁,54, the information contained in subcounter C7 ismoved to a FIFO register. Again, the reason for this is that because ofthe geometry of the preferred system, when the electron beam ispositioned behind the next selected aperture AP₅₁,54, no x-raymicro-beam will pass through IP₁ and strike any logical detector in thearray. In accordance with this embodiment, no x-ray micro-beam will passthrough IP₁ and strike any logical detector until the electron beam ispositioned behind AP₅₂,50.

As the electron beam is being positioned behind AP₅₂,50, subcounter M4is loaded with the information stored in the FIFO corresponding to thepartial pixel information output from subcounter C7. While the electronbeam is positioned behind AP₅₂,50 the partial image pixel informationdetected by logical detector M4 is input to and added to the contents ofsubcounter M4.

As the electron beam is positioned behind the next aperture in the samecollimator row AP₅₂,51, the information contained in subcounter M4 isinput to subcounter J4. While the electron beam is positioned behind thenext selected aperture AP₅₂,51 an x-ray micro-beam will pass through IP₁and strike logical detector J4. The partial image pixel informationdetected by logical detector J4 will be input to and added to thecontents of subcounter J4.

As the electron beam is positioned behind the next aperture in the samecollimator row AP₅₂,52, the information contained in subcounter J4 ismoved to subcounter F4. While the electron beam is positioned behind thenext selected aperture AP₅₂,52, an x-ray micro-beam will pass throughIP₁ and strike logical detector F4. The partial image pixel informationdetected by logical detector F4 will be input to and added to thecontents of subcounter F4.

As the electron beam is positioned behind the next aperture in the samecollimator row AP₅₂,53, the information contained in subcounter F4 ismoved to subcounter C4. While the electron beam is positioned behind thenext selected aperture AP₅₂,53, an x-ray micro-beam will pass throughIP₁ and strike logical detector C4. The partial image pixel informationdetected by logical detector C4 will be input to and added to thecontents of subcounter C4.

As the electron beam is positioned behind the next aperture in the samecollimator row AP₅₂,54, the information contained in subcounter C4 ismoved to a FIFO register. Again, the reason for this is that because ofthe geometry of the preferred system, when the electron beam ispositioned behind the next selected aperture AP₅₂,54, no x-raymicro-beam will pass through IP₁ and strike any logical detector in thearray. In accordance with this embodiment, no x-ray micro-beam will passthrough IP₁ and strike any logical detector until the electron beam ispositioned behind AP₅₃,50.

As the electron beam is being positioned behind AP₅₃,50, subcounter M1is loaded with the information stored in the FIFO corresponding to thepartial pixel information output from subcounter C4. While the electronbeam is positioned behind AP₅₃,50 the partial image pixel informationdetected by logical detector M1 is input to and added to the contents ofsubcounter M1.

As the electron beam is positioned behind the next aperture in the samecollimator row AP₅₃,51, the information contained in subcounter M1 ismoved to subcounter J1. While the electron beam is positioned behind thenext selected aperture AP₅₃,51 an x-ray micro-beam will pass through IP₁and strike logical detector J1. The partial image pixel informationdetected by logical detector J1 will be input to and added to thecontents of subcounter J1.

As the electron beam is positioned behind the next aperture in the samecollimator row AP₅₃,52, the information contained in subcounter J1 ismoved to subcounter F1. While the electron beam is positioned behind thenext selected aperture AP₅₃,52, an x-ray micro-beam will pass throughIP₁ and strike logical detector F1. The partial image pixel informationdetected by logical detector F1 will be input to and added to thecontents of subcounter F1.

As the electron beam is positioned behind the next aperture in the samecollimator row AP₅₃,53, the information contained in subcounter F1 ismoved to subcounter C1. While the electron beam is positioned behind thenext selected aperture AP₅₃,53, an x-ray micro-beam will pass throughIP₁ and strike logical detector C1. The partial image pixel informationdetected by logical detector C1 will be input to and added to thecontents of subcounter C1.

As the electron beam is positioned behind the next aperture in the samecollimator row AP₅₃,54, the information contained in subcounter C1 isoutput to a monitor where the information can be Processed for display.The output of C1 will contain complete image pixel information for pixelIP₁. The reason for this is that because of the geometry of thepreferred system, no other apertures, other than those identified abovewill include an x-ray micro-beam that passes through IP₁.

In accordance with the string method, at the same time that informationfor IP₁ is being collected in string 1, information for IP₂ is beingcollected in string 2. The complete image pixel information for IP₂ iscollected in accordance with the same method as with string 1 exceptthat the logical detectors for string 2 (L10, H10, E10, B10, L7, H7, E7,B7, L4, H4, E4, B4, L1, H1, E1 and B1) collect the information andsubcounters L10, H10, E10, B10, L7, H7, E7, B7, L4, H4, E4, B4, L1, H1,E1 and H1 combine the information as the electron beam is positionedbehind AP₅₀,50, AP₅₀,51, etc. The same process is also used with thecorresponding subcounters and logical detectors for strings 3, 4, 5, 6,7, 8 and 9. Thus, after the electron beam has been positioned behindaperture AP₅₃,53 complete information for nine image pixels can beoutput to a video monitor for display. Also, a new string 1, string 2, .. . , and string 9 are started every time the electron beam ispositioned behind a new aperture. It should be noted that a FIFO may bereplaced by any storage mechanism that can store the intermediateoutputs of the strings until complete image pixel information has beencollected.

As shown in FIG. 36, preferably only 96 of the 144 logical detectors areactive. For example, string 1 is preferably mapped to 11 active (J10,F10, C10, M7, J7, F7, C7, J4, F4, C4 and F1) and 5 inactive detectorelements (M10, M4, M1, J1 and C1). Since these inactive detectorelements do not provide any information about the image pixel a zerovalue is input into the corresponding subcounters. Also, as noted, if itis determined that the inactive logical detectors will never be used tocollect image pixel information, they need not have an actual detectorassociated with them. Similarly, for string 2, the inactive logicaldetectors (those outside the active detector array area 822) also inputzero values to their corresponding subcounters, and so forth with theother strings.

When less than all of the logical detectors are active it is preferableto normalize the complete image pixel information to account fordifferences in the number of active detector elements for each string.As shown in Table II, the number of active detector elements providinginput data vary between 10, 11 or 12 depending on the particular string.The complete image pixel information from each of the nine strings ispreferably normalized by dividing the complete image pixel informationby the number of active detectors in that string and then multiplying by12.

                  TABLE II    ______________________________________    STRING    NUMBER OF ACTTVE ELEMENTS    ______________________________________    1         11    2         10    3         11    4         10    5         12    6         10    7         11    8         10    9         11    ______________________________________

Alignment

The electron beam 40 in x-ray source 10 is preferably precisely alignedsuch that it will illuminate the area on the target layer at the exactpoint at which the axis of the collimator hole intersects the targetlayer. When no object is interposed between the target 50 and themulti-detector array 110, such a precisely aligned electron beam willresult in a near symmetrical distribution of x-ray intensity across theface of detector elements 160 of the multi-detector array 110, withinthe pseudo-circle 400. An electron beam which is not so preciselyaligned may create a non-symmetrical distribution of x-ray intensitiesacross the face of the detector elements 160 of the multi-detector array110.

Alignment of the electron beam behind the collimator holes is preferablyaccomplished with a 2-step process. An initial alignment procedure ispreferably performed to approximate the correct positioning of theelectron beam 40. The initial alignment procedure is preferably followedby a fine alignment procedure that optimizes the position of the centerof the electron beam profile relative to the collimator holes.

The first step of the preferred initial alignment procedure is comprisedof locating the electron beams using a-priori knowledge related to thephysical, electrical and magnetic properties of the scanning system. Therelative spacing of the electron beam positions may be reasonablycorrect at this point, but the absolute positions of the electron beamsmay not be because of the difficulty in indexing the electron beamposition array to the collimator holes, and because of small cumulativeerrors. Therefore, a "dithering" process is preferably employed wherebyseveral measurements are made by making small adjustments of the indexposition for a whole array of electron beam positions. Typically, 25measurements are made where the index point is moved in a 5 by 5 x-ygrid. The total size of the grid is approximately the spacing of onecollimator hole. The data collected for each measurement consists of thetotal intensity measured by the multi-detector array for each of thecollimator holes.

The collected data for a give collimator hole will preferably be anarray of 25 values. Many of the values will indicate that little or nox-ray flux impinged upon the multi-detector array, but several willindicate that at least part of the electron beam produced x-rays thatimpinge upon the multi-detector array. An approximate optimum beamposition location is determined by mathematically fitting amulti-dimensional surface to the illuminated data.

Thus, approximate optimum beam positions are determined by thisprocedure. These positions are refined using the fine alignment methoddescribed below.

To initiate the fine alignment procedure, initial x-deflection valuesand y-deflection values are preferably computed for each collimatoraperture, employing the initial alignment procedure described above.Using these computed initial deflection values the electron beam isscanned across the target, momentarily stopping at each of the computedlocations corresponding to the computed x and y deflection values. Thepartial image pixel information obtained from each detector element foreach x-ray pencil beam generated by stopping at each computed locationis analyzed for even distribution of the x-ray intensity over severalframes. (A complete scan of the target is referred to as a frame.) Ifthe analysis results in a determination that the distribution of x-rayintensity is not even, new x-deflection values and/or y-deflectionvalues are calculated and the alignment procedure is repeated to ensureoptimal distribution.

The preferred way of analyzing the distribution of x-ray flux across theface of the multi-detector array is to compare the average intensity ofthe x-ray rays detected by selected areas of the face of themulti-detector array. This is preferably accomplished by dividing thepreferred ninety-six detector element multi-detector array into eightareas comprising substantially the same number of detector elements.

FIG. 38 is a representational diagram of the face of a multi-detectorarray divided into eight areas. Each of the eight areas is referred toas an "octant." The eight octants are identified as the top right outeroctant ("TRO") 1345, top right inner octant ("TRI") 1346, top left outeroctant ("TLO") 1349, top left inner octant ("TLI") 1350, bottom rightinner octant ("BRI") 1347, bottom right outer octant ("BRO") 1348,bottom left outer octant ("BLO") 1351, and bottom left inner octant("BLI") 1352.

In the preferred embodiment, the 96 detector elements 1339 are evenlydivided among the eight octants. Therefore each octant contains 12detector elements 1339. However, it is contemplated that otherarrangements may be used in the present invention. For example, analternative arrangement could consist of 13 detector elements 1339associated with each inner octant, with 11 detector elements 1339associated with each outer octant.

The beam alignment calculations are preferably determined separately forthe x-axis 1662 and the y-axis 1660. The preferred sequence of steps todetermine the proper beam alignment along the y-axis 1660 are asfollows. The process begins when an x-ray pencil beam from a singlecollimator aperture strikes the scintillator elements of multi-detectorarray 110.

The total intensity values for each octant is summed by counting thenumber of x-ray photons which are received by each detector element 1339associated with each octant. For example, arbitrarily selecting thevariable V to refer to the sum of the photon counts in a particulararea, V_(TRO) is the sum of all the photon counts in the TLO octant1345. Similarly, V_(TRI) is the sum for the TRI octant 1346, V_(TRO) forTLO octant 1349, V_(TLI) for TLI octant 1350, V_(BLO) for BLO octant1351, V_(BLI) for BLI octant 1352, V_(BRI) for BRI octant 1347, andV_(BLO) for BLO octant 1348. The intensity values for each octant, foreach x-ray pencil beam from each collimator aperture for each of apredetermined number of succeeding frames, is accumulated. The presentlypreferred embodiment uses the octant values from 100-120 frames toperform the beam calculations. Thus, there are a total of eight octantvalues for each beam/aperture combination.

The accumulated values for the octants in the top and bottom halves ofthe PMT array are then separately summed. Thus the top octantaccumulated value is V_(top) =V_(TRO) +V_(TRI) +V_(TLI) +V_(TLO). Thebottom octant accumulated is V_(bottom) =V_(BRO) +V_(BRI) +V_(BLI)+V_(BLO).

Next the top octant accumulated value is compared to the bottom octantaccumulated value. This comparison produces a y-axis alignment factor(AF_(y-axis)) which is a measure of the accuracy of the x-ray beamalignment with respect to a particular aperture along the y-axis. Theformula to determine the AF_(y-axis) is: ##EQU7##

If the electron beam is properly aligned with the aperture underanalysis along the y-axis, the accumulated intensity values for the topand the bottom octants should be the same. Thus when V_(top) is equal toV_(bottom), AF_(y-axis) =0 and the beam is properly aligned along they-axis for the aperture under analysis.

If the electron beam was positioned to favor the top half of themulti-detector array, then V_(top) will be greater than V_(bottom). Thisresults in AF_(y-axis) >0. If the electron beam is positioned to favorthe bottom half of the multi-detector array, then V_(top) will be lessthan V_(bottom). This results in AF_(y-axis) <0. The value ofAF_(y-axis) generally indicates the amount the y-deflection value shouldchange to optimize the alignment.

The method to determine the optimal electron beam alignment along thex-axis is similar. For this calculation, the accumulated values for theleft and right octants are separately summed. Thus the right octantaccumulated value is V_(right) =V_(TRO) +V_(TRI) +V_(BRO) +V_(BRI). Theleft octant accumulated value is V_(left) =V_(TLO) +V_(TLI) +V_(BLO)+V_(BLI). The formula to determine the x-axis alignment factor(AF_(x-axis)) is: ##EQU8## Calculations almost identical to those usedfor the y-axis alignment are used to determine the optimal alignment ofthe electron beam along the x-axis.

The x-axis and y-axis alignment factors are transmitted to the controlcomputer 890. Control computer 890 processes these alignment factors todetermine the amount of correction required at the x-ray source 798 tooptimally align the x-ray pencil beam. The control computer 890 nextupdates the beam deflection lookup tables 918.

By adjusting the electron beam's positioning on the target 1250, x-raysare emitted from the target at a different position relative to thecollimator grid aperture. The x-ray pencil beams passing through thecollimator grid aperture would then illuminate the multi-detector arrayat a corrected optimally aligned position.

This alignment may be performed whenever the system is activated, atpreset intervals or continuously.

While the previous discussion explores alignment calculations along thex and y axes, other octant calculation methods are also contemplatedwithin the boundaries of this aspect of invention. For example, angularalignment calculations may be performed by comparing the accumulatedvalue of the top right octants with the values of the bottom leftoctants and the too left octants with the bottom right octants.

Beam Alignment Extractor and Image Reconstruction Engine

FIG. 39 is a block diagram of the circuitry for the beam alignmentextractor and image reconstruction engine. The output signals of thesignal conditioner 810 are input, via connectors 1655, to the beamalignment extractor.

The conditioned partial image pixel signals from each detector elementfor each step of the electron beam are input to RTE octant counters1354. There are preferably eight RTE octant counters. Each of the octantcounters receives the conditioned partial image pixel signals from oneof the eight octants. Each RTE octane counter 1354 splits theconditioned partial image pixel signals into two essentially identicalsignals. One set of the conditioned partial image pixel signals is usedto analyze the optimal alignment of the electron beam. The other set istransmitted to the image reconstruction engine ultimately to be used toreconstruct the image of the object being investigated.

Each RTE octant counter 1354 then processes the input conditioned imagepixel signals to obtain a total photon count for its correspondingoctant. In sequential order, the RTE octant counters 1354 next transmitthe total photon sum for each octant to the frame-summation chip 1357.This process is controlled and outputs to the frame summation chip areenabled by control signals communicated to the RTE octant counters 1354from the gain & alignment engine 1674.

The frame-summation chip 1357 is an arithmetic logic unit ("ALU"). Foreach photon count input from a RTE octant counter 1354, theframe-summation chip 1357 also inputs an accumulated octant value fromthe gain & alignment memory 1678. This accumulated octant valuecorresponds to the sum of the photon counts from one or more previousframes for the same octant and for the same aperture on the collimatorwhich was illuminated to produce the present photon count. The framesummation chip 1357 adds the photon count to the accumulated octantvalue to produce a new octant value, which is then stored at the gain &alignment memory 1678.

The gain & alignment engine 1674 controls the operation of the octantcounters 1354, frame-summation chip 1357 and gain & alignment memory1678. After approximately 100-120 frames of information have beencollected at the gain & alignment memory 1678, the gain & alignmentengine 1674 communicates instructions to the gain & alignment memory1678 to output the beam alignment information, which is transmittedthrough transceivers 1704 and 1706 to the RTE output circuits 818.

The string counters 1372 input partial image pixel signals from the RTEoctant counters 1354. The string counters 1372 process partial imagepixel values to reconstruct data values for complete image pixels, asexplained more fully in the detailed description of FIG. 37.

During the image reconstruction process, partially constructed imagepixel values are stored by the string counters 1372 at the line FIFO("first in first out") chips 1702. After 166 items of partial imagepixel values are input into a line FIFO chip 1702, each successive itemstored at that line FIFO chip 1702 will cause the line FIFO chip 1702 totransmit the then earliest stored string data value back to the stringcounters 1372. Line FIFO 1702 receives timing signals from the RTEcontrol chips.

The string counters 1372 transmit data values for complete image pixelsto the normalization PROM 1692. The normalization PROM adjusts this datavalue based upon the number of active detector elements which contributepartial image pixel information for that image pixel (this is explainedin more detail in conjunction with the detailed discussion of FIG. 37).The normalization PROM 1692 receives control signals from the imagereconstruction controller 1696 through electrical connection 1768.

Normalization PROM 1692 outputs normalized image pixel information tothe output FIFOs 1700 through electrical connection 1746. Three lines ofnormalized image pixel information are stored at the output FIFOs 1700before the normalized image pixel data is transmitted to the RTE outputcircuit.

The normalized image pixel information from the normalization PROM 1692is also input to the image memory unit 1694. The normalized image pixelinformation for the entire image is stored and properly ordered at theimage memory unit 1694. The control computer can access this image datathrough transceiver 1698.

The image reconstruction controller transmits the control signals whichoperate the components of the image reconstruction engine. Control andaddressing signals are communicated to the image memory unit 1694 onelectrical connections 1758 and 1756. Control signals are sent, viaelectrical connection 1768, to the normalization PROM. The imagereconstruction controller 1696 communicates control signals, viaelectrical connection 1724, to the string counters 1372.

Control information from the control computer 890 are input to thereal-time eye through RTE input circuit 1620, which receives lightpulses from high-speed fiber-optic cable 824. The RTE input circuit 1620comprises a light detector and circuitry which detects and demodulatesthe light pulses into electrical signals which contain the controlinformation from the control computer 890. The control information issent from RTE input circuit 1620 to the RTE control chips 1690 throughelectrical connection 1714.

The RTE control chips 1690 send timing signals to the RTE circuitrythrough electrical bus connection 1710. The RTE control chips 1690 sendcontrol signals to the RTE circuitry through electrical bus connection1760.

RTE output circuit 818 sends image reconstruction and gain & alignmentinformation to the control computer 890 through high-speed fiber-opticcable 826. RTE output circuit 818 comprises a high radiance LED andcircuitry which converts electrical signals into light pulses.

Turning to FIGS. 40A-E, a detailed diagram is presented of the RTE inputcircuit 1620. RTE input circuit 1620 receives light pulses from righttransmitter 880. Light pulses are detected and converted to anelectrical signal by the fiber-optic receiver 1612. The electricalsignal is filtered and shaped by circuits 1622 and 1624. The electricalsignal is then input to the taxi chip 1614, a standard AM7969 chipavailable from AMD Corp, which functions as a serial to parallelconverter. The electrical input signal was necessarily in a serialformat because of its transmission through a fiber-optic cable. Fourbits of control signals and eight bits of data signals are output fromtaxi chip 1614. While the present description of FIGS. 40A-E is directedto the RTE input circuit 1620 of the data receiver 812, a similarcircuit exists for other components of the present invention whichreceives light pulses through fiber-optic cables.

Phase locked loop (PLL) circuit 1629, located in data receiver 812,receives and locks onto a master 12.5 Mhz clock signal that is generatedin the programmable scan controller 920 (FIG. 22). This master clocksignal drives both the taxi chip 1614 in the data receiver 812 and thetaxi chip 1602 in the data transmitter 818 to generate an output at aclock rate of 12.5 Mhz. MC88915 clock doublers 1628, 1630, and 1632 areused to quadruple the 12.5 Mhz clock signal to a 50 Mhz frequency.Timing circuit 1626 uses this 50 Mhz clock to synchronize taxi chip 1614with the other components of the beam alignment extractor and imagereconstruction circuitry. Timing circuit 1626 generates a data strobesignal which is transmitted via electrical connection 1636. Timingcircuit 1626 generates a control strobe signal which is transmitted viaelectrical connection 1634.

FIGS. 41A-E diagram the RTE output circuit 818, which is also referredto as the right data transmitter. Taxi chip 1602 is another standardAM7968 chip available from AMD Corp. which also functions as a parallelto serial converter. Parallel data bits from the image reconstructionengine and the beam alignment extractor are input to the taxi chip 1602,which outputs a serial data signal. This serial data signal is thenshaped by the conditioning circuitry 1610. The output signal from theconditioning circuit 1610 is sent to fiber-optic transmitter 1604, whichtransforms the serial data signal into light pulses through the use of ahigh-radiance LED. The light pulses are sent to a data receiver 880through high-speed fiber-optic cable 826. While the present descriptionof FIGS. 41A-B are directed to the RTE output circuit 1620 of thereal-time eye, a similar circuit exists for other components of thepresent invention which transmits light pulses through fiber-opticcables.

FIGS. 42A-E are a circuit diagram of the RTE control chips 1690 whichare located in the data receiver 812. Information from the controlcomputer 890 that is acquired through the RTE input circuit 1620 isdistributed to the various components of the multi-detector arraythrough the RTE control chips 1690, each of which is a MACH435programmable IC chip available from AMD Corp. Data outputs from the RTEinput taxi chip 1614 are input to the RTE control chips via 8 bitelectrical connection 1616. Control information outputs from the RTEinput taxi chip 1614 are input to the RTE control chips via 4 bitelectrical connection 1616.

Data acquisition control chip 1638 distributes control informationrelating to the selection of data that is acquired and processed by thecomponents of the multi-detector array 822. Host memory control chip1642 communicates instructions to the image memory unit 1694 and thegain & alignment memory unit 1678. Timing control chip 1640 communicatestiming and diagnostic signals to the circuitry of the beam alignmentextractor and the image reconstruction engine. The timing controlsignals for the signal conditioner 1510 (FIGS. 33A-I) is output from thetiming control chip 1640 through connection 1646. 1 Kbyte of nonvolatilememory 1644 stores calibration information for the circuitry of the beamalignment extractor and the image reconstruction engine.

Referring to FIGS. 43A-B four connectors 1650, 1652, 1654, and 1656 formthe sensor connections 1655 between the signal conditioner 810 and theRTE octant counters 1354. After signal conditioning, signals from eachof the 96 PMT detector elements 1339 connects the RTE octant counters1354 through one of 96 electrical connections on the four connectors1650, 1652, 1654, and 1656.

FIGS. 44A-B are diagrams of the preferred octant counters 1354. Eightsuch octant counters 1354 are used in the real-time eye. Each octantcounter 1354 preferably comprises an ISP1032TQ lattice IC chip. Octantcounter 1303A processes the inputs from the photocathode elements 1339which is associated with the TLO octant. Similarly, octant counter 1303Bprocesses the inputs for the TRO octant, octant counter 1303C for theTLI octant, octant counter 1303D for the TRI octant, octant counter1303E for the BLO octant, octant counter 1303F for the BRO octant,octant counter 1303G for the BLI octant, and octant counter 1303H forthe BRI octant.

Each octant counter 1354 contains data input connections for each of the12 PMT photo-cathode elements 1339 that is preferably associated witheach octant. Upon detection of light photons by a PMT photo-cathodeelement 1339, an electrical signal is sent to its corresponding octantcounter 1354. For the x-ray pencil beam which passes through a singlecollimator aperture, each of the eight octant counters 1354 produces a9-bit value which contains the intensity data from all 12 of eachoctant's associated PMT photo-cathode elements 1339.

FIG. 45 diagrams the frame-summation chip 1357, which is an arithmeticlogic unit ("ALU") and is preferably a L4C381JC26A IC chip availablefrom Logic Devices, Inc. The 9 bit output from each octant counter 1354is input to the frame-summation chip 1357 through connection 1356. Theframe-summation chip 1357 processes eight numbers for each collimatoraperture. For each succeeding frame, the frame-summation chip 1357 sumsthe corresponding values for the same octant for the same aperture fromthe previous frames. In the preferred embodiment, the octant values for100-120 frames are added together to construct the data used for x-raybeam alignment.

FIG. 46 diagrams the gain & alignment engine 1674, which is preferably aMACH435 IC chip available from AMD Corp. The gain & alignment engine1674 determines the items of beam alignment data which is to beprocessed and manner of processing intended for that item of data.Additionally, the gain & alignment engine 1674 controls the timing ofthe components within the beam alignment extractor 816.

FIG. 47A-D are a diagram showing the gain & alignment memory chips 1678which is comprised of four 1-Mbyte SRAM memory chips 1664, 1666, 1668,and 1670, each available under the model number MM624256AJP-20 fromHitachi Corporation. For each frame, eight values are collected for eachcollimator aperture. These values are stored in the gain & alignmentmemory 1678 after being processed through the frame-summation chip 1357.After 100-120 frames, the control computer will preferably access andprocess the data which is stored in the gain & alignment memory 1678 tocorrect the alignment of the x-ray beam.

FIGS. 48A-I are diagrams of string counters 1372 for strings one throughnine. Nine such string counters are used in the image reconstructioncircuitry, each of which is preferably comprised of two gate arrays 1680and 1682, preferably IC part numbers ISP1032TQ available from LatticeCorp. Each string counter 1372 contains a total of 16 subcounters, witheach gate array 1680 and 1682 containing 8 individual subcounters. Sincethere are a total of nine string counters 1372 and each string counter1372 contains 8 individual subcounters, the total number of individualsubcounters is 144. Since all nine string counters 1372 functionsimilarly, only the string one string counter 47A will be discussed indetail.

Referring back to FIG. 36, the detector element inputs for string oneare marked with "1." Since only the 96 elements within the center pseudocircle 1780 correspond to active detector elements, the inputs mappedfrom active detector elements for string one are element F1, J4, F4, C4,M7, J7, F7, C7, J10, F10, and C10. Referring to FIG. 48A, dataconnection 1782 on string counter 1372 is mapped to the input fromactive detector element F1. Data connections 1784, 1786, 1788, 1790,1792, 1794, 1796, 1798, 1800, and 1802 are mapped to active detectorelements J4, F4, C4, M7, J7, F7, C7, J10, F10, C10 respectively. Theinputs from elements M1, J1, C1, M4, and M10 correspond to the inactivedetector elements for string counter one. Therefore, data connections1804, 1806, 1808, 1810, and 1812 which correspond to inactive detectorelements M1, J1, C1, M4, and M10 respectively, are all tied to ground.

Each of the other string counters are similarly mapped to theirrespective active and inactive detector elements. FIG. 36 diagrams the144 logical detector elements and the string counter number that theyare mapped.

FIG. 48A also diagrams a line FIFO ("first in first out") chip 1702,preferably part no. CY7C441-14VC available from Cypress Semiconductor.The RTE image reconstruction circuitry consists of nine line FIFOs, witheach line FIFO 1702 paired with a corresponding string counter 1372.Partial image pixel values from string counter 1372 are input into lineFIFO 1702 as the string counter acquires data from the multi-detectorarray. After the 166th string counter data value is input into the lineFIFO 1702, every succeeding data value will cause the line FIFO 1702 toreturn a data value that was stored 166 counts before to the itscorresponding string counter 1372. The data values are returned to thestring counter 1372 to be summed with new data values acquired for thesame image pixel.

FIG. 49 diagrams the normalization PROM 1692, which is a CY7C235A-18JCCypress Semiconductor unit. As explained previously, each string counter1372 is mapped to 16 logical detector elements on a 12-by-12 array oflogical detector elements. However, each string counter 1372 may receivemeaningful input data from only among the 96 active detector elementswhich comprise the center pseudo-circular array. As shown in FIG. 36,the number of active detector elements providing input data may be 10,11 or 12 depending on the particular string counter. The normalizationPROM 1692 normalizes the outputs from the nine string counters 1372 bycalculating the proper output levels based upon the number of activeinput detector elements for each string.

FIG. 50 diagrams the image reconstruction controller 1696. Imagereconstruction controller 1696 functions as the "engine" within theimage reconstruction engine 814. Image reconstruction controller 1696controls the timing and operation of the string counters 1372. The imagereconstruction controller 1696 is the component which keeps track ofboth the strings and the individual image pixels which arereconstructed. Image reconstruction controller 1696 also controls theoperation of the image memory unit 1694.

FIG. 47A-D diagram the image memory unit 1694, which is preferablycomprised of two conventional 1-Mbyte MM624256AJP-20 SRAM chips 1816 and1818. The normalized pixel data output from the normalization PROM isinput to the image memory unit 1694 through electrical connection 1718.The image memory unit 1694 combines and correctly orders the pixel datafor the entire image.

FIG. 51A-C are a diagram of the output FIFOs 1700. The output FIFOs 1700preferably comprises three CY7C453-14JC devices 1820, 1822, and 1824.The output FIFOs 1700 store three lines of pixel data before outputtingthis data in frame order through RTE output circuit 818. The outputFIFOs 1700 function in this manner because a completed scan of onecollimator aperture row will result in the completion of three lines ofimage pixels. The pixel data is continually input from the normalizationPROM 1692 until the three lines of pixel data are stored.

FIG. 52 diagrams the output FIFO controller 1814. Output FIFO controller1814 consists of a lattice chip available under the model number MACH435IC chip from AMD Corp. Output FIFO controller 1814 controls theoperation of the three output FIFOs 1700.

Detector Controller

The detector controller 805 (FIG. 21) for the detectors 822 and 1522receives image pixel data and beam alignment data from the detectors andtransmits control information to the detectors. Right receiver 880optically receives image pixel data and beam alignment data from theright detector 822 through high-speed fiber-optic cable 826.Consequently, right receiver 880 includes the light signal to electricalsignal conversion circuitry described more fully in conjunction withFIGS. 40A-E. The left receiver 846 operates in a similar fashion toreceive image pixel data and beam alignment data from the left detector1522.

Right transmitter 886 optically transmits threshold and gain controldata to the right data receiver 812 though fiber-optic cable 824.Consequently, beam controller interface 794 includes the electrical tolight conversion circuitry described more fully in conjunction withFIGS. 41A-B. Right transmitter 886 also receives synchronization signalsfrom beam deflection lookup table 918 (FIG. 25). The left transmitter848 operates in a similar manner to communicate control signals to theleft detector 1522.

FIG. 53A-E are a circuit diagram of the preferred control logic for thedetector controller 805. Control PAL 1870, preferably a conventionalMACH435 programmable IC chip, provides the control signals to coordinatethe activity of the detector controller 805 and the left and rightmultiplexer PALs 1872 and 1874. Control PAL 1870 receives left datacontrol signals LRC from the left receiver 846 via leads 1875.Similarly, control PAL 1870 receives right data control signals from theright receiver via leads 1877. Control PAL 1870 receives timing inputcontrol signals through a tube controller link 1876.

Left multiplexer PAL 1872 preferably functions as data multiplexers forthe items of data which are received and transmitted by the leftreceiver 846 and the left transmitter 848. Left multiplexer PAL 1872preferably loads gain and threshold control data from a bidirectionaldata bus via leads 1873. Left multiplexer PAL 1872 also preferably loadsimage pixel data and alignment data from the left receiver 846 via leads1877.

Left multiplexer preferably transmits gain and threshold controlsignals, via leads 1879, to the left transmitter 848, which thenoptically transmits this data to the left detector 1522. The alignmentdata received by the left multiplexer PAL 1872 is redirected to thecontrol computer via the bidirectional data bus via leads 1873. Theimage pixel data received by the left multiplexer PAL 1872 issequentially redirected to a data FIFO 1878. Data FIFO 1878 andconnector 1880 function as a data interface between the left multiplexerPAL 1872 and the left frame buffer 850. Right multiplexer PAL 1874functions similarly to control and select the items of data which arereceived and transmitted by the right receiver 880 and right transmitter886. Similarly, data FIFO 1882 and connector 1884 function as the datainterface between the right multiplexer PAL 1874 and the right framebuffer 872.

In the preferred embodiment, detector controller 805 is fabricated as aPC module that plugs into the bus of the control computer 890. Thepreferred interface circuitry between detector controller 805 and thecontrol computer 890 is described more fully in connection with thedetailed description of FIG. 54B.

Tube Controller

Tube controller 807 generates scan control data which directs theoperation of the beam controller 796, thereby controlling the scanningpattern of the x-ray source 798. Tube controller 807 functionallycomprises a beam deflection lookup table 918 which stores beamdeflection data for each point on the target anode, programmable scancontroller 920, beam transmitter 916, I/O transceiver 964, and I/O latch958.

FIG. 54A is a circuit diagram of the control logic for the tubecontroller 807. Control PAL 1402 generally performs the functions of theprogrammable scan controller 920, by processing control instructionsreceived from control computer 890, and distributing signals forloading, running or stopping the scan patterns to the memory control PAL1404 and data PAL 1406. Control PAL 1402 provides control signals todirect the operation of the components within the tube controller 807.For example, control PAL 1402 is programmed to set the "measure andmove" times for each collimator hole scanned. The preferred programmingmodule for the control PAL 1402 is included in Appendix A.

Beam deflection lookup table 918 is preferably comprised of a memorycontrol PAL 1404 and lookup table memory 1408. Memory control PAL 1404generates control signals to direct the storage and retrieval ofinformation in the lookup table memory 1408. At appropriate times,memory control PAL 1404 directs the retrieval of the beam deflectiondata from the lookup table memory 1408. The retrieved beam deflectiondata is sent to the beam transmitter 916 for transmission to the beamcontroller 796. Beam transmitter 916 is preferably a conventionaloptical transmission circuit, which is discussed more fully inconnection with the detailed description of FIGS. 41A-B.

Data PAL 1406 is programmed to function as the data multiplexer for thedata which is received or transmitted by the tube controller 807.

In the preferred embodiment, tube controller 807 is fabricated as a PCmodule that plugs into the bus of the control computer 890. FIG. 54B isa diagram of the preferred interface circuitry for connecting the tubecontroller 807 and the control computer 890. Connectors 1425 and 1427interface the tube controller 807 with the bus of the control computer890. Data transceivers 1401 and 1403 transfer binary information betweenthe tube controller 807 and the control computer 890 on a three statebidirectional 16 bit data bus B 0..15! via leads 1409. I/O read controlsignals are applied via lead 1405 and I/O write control signals areapplied via lead 1407.

The tube controller also controls the I/O transceiver 964 and the I/Olatch 958. The tube controller 807 directs the various sets of controlsignals received from the control computer to the I/O transceiver 964and the I/O latch 958 for further transmission of these control signalsover an optical link to the I/O controller. I/O transceiver 964preferably includes optical communications circuitry which is more fullydiscussed in connection with the detailed description of FIGS. 40A-E and41A-B.

Beam Controller

FIGS. 55A-E diagram the control logic within the beam controllerinterface 794, which processes and distributes analog coil currentcontrol signals to the various coil drivers. The digital scan controldata generated by the tube controller 807 is optically coupled to thebeam controller input circuit 1408, which preferably includes theoptical communications circuit described more fully in connection withthe detailed description of FIGS. 40A-E. Beam controller input circuit1408 outputs eight parallel bits of digital scan control data to aneight-bit data bus D 0..7! and four parallel bits of control data CD toa control PAL 1410, which distributes and/or reformats the digital scancontrol data within the beam controller interface 794.

Referring to FIGS. 55B-C, control PAL 1410 preferably outputs controlsignals, via leads 1411 (LD1 and LD2), to instruct the x-deflection PAL1412 to sequentially load parallel bits of digital x-deflection coilcontrol data DXDEF from the eight-bit data bus D 0..7!. The x-deflectionPAL 1412 essentially manipulates the digital x-deflection coil controldata DXDEF to generate a smoothly ramping triangular waveform at thex-deflection driver 778. Approximately every 1.28 μsec, the x-deflectionPAL 1412 preferably converts the parallel bits of digital x-deflectioncoil control data DXDEF to serial bits of digital x-deflection coilcontrol data SDX. The serial x-deflection coil control data SDX iscoupled, via output line 1413, to a twenty-bit serial DAC 1414 whichconverts the information to an analog signal that is preferably appliedto an intermediate x-deflection amplifier 1416.

Approximately every 80 nsec, the x-deflection PAL 1412 mathematicallymanipulates the sequentially acquired items of digital x-deflection coilcontrol data DXDEF to calculate an eight-bit x-slope value, which isreferred to as the x-slope control data XSD. The x-slope control dataXSD is transmitted to DAC 1418 for conversion to an analog signal, andits analog output signal is preferably coupled to a series ofintermediate x-slope amplifiers 1420. The amplified analog x-slopecontrol signals XSD is preferably summed with the amplified analogx-deflection coil control data SDX to generate a smoothly ramping outputwaveform, which is amplified by intermediate amplifier 1417 to producethe x-deflection coil control signal XDEFL. The x-deflection coilcontrol signal XDEFL is preferably output, via output line 1418, to apreferred x-deflection driver 778, which is described more fully inconnection with the detailed description of FIG. 56. Alternatively, thex-deflection coil control signal XDEFL can be coupled, through anamplifier 1419 and a BNC connector 1444, to a commercially availableamplifier, for example a Centronics amplifier, which then drives thecurrent in the x-deflection coil.

Analog y-deflection coil control signals are generated in the samefashion and output to a y-deflection driver 782. However, if a rasterscan pattern is employed, then the serial y-deflection coil control dataSDY is directly generated by the control PAL 1410, therefore ay-deflection PAL, y-slope control data YSD, and related circuitry arenot required.

Control PAL 1410 also outputs control signals, via leads 1421 (LD6, LD7,LD8, LD9, and LD10), to instruct the small DAC control PAL 1422 tosequentially load x-step control data (XCD), dynamic focus coil controldata (DFCD), and stigmator control data (SCD) from the data bus D 0..7!.Small DAC control PAL 1422 redistributes the XCD and DFCD controlsignals to multi-channel DAC 1426 and redistributes SCD control signalsto multi-channel DAC 1424. DAC 1424 preferably outputs analog 0°stigmator coil control signals to the 0° stigmator driver 786 through anintermediate 0° amplifier 1428. Analog 45° stigmator coil controlsignals are similarly output to the 45° stigmator driver through anintermediate 0° amplifier 1430. DAC 1426 preferably outputs analogx-step slope control signals XSTEPSLP to the x-step driver 780 viaoutput line 1432. Similarly, analog x-step amplitude control signalsXSTEPAMP are preferably output to the x-step driver 780 via output line1434 and analog dynamic focus coil control signals DFOCUS are preferablyoutput to the dynamic focus driver 776 via output line 1436.

Serial data PAL 1438 preferably receives static focus coil control dataSDIN from the I/O controller 762. Serial data PAL 1438 couples controldata SDIN to a DAC 1440, which converts this information to analogstatic focus coil control signals which are sent to the static focusdriver 774 through intermediate focus amplifiers 1442.

The analog coil control signals from the beam controller interface 794are preferably transmitted to suitable power amplifier circuits withinthe coil drivers to drive the current patterns in their correspondingfocus or deflection coils. For example, the analog x-deflection coilcontrol signals XDEFL from the beam controller interface 794 arepreferably coupled, via input line 1418, to a preferred x-deflectiondriver 778 (FIG. 56). The XDEFL control signals are applied to a control15 amplifier 1454, which regulates the activity of power amplifiers 1446and 1448. The x-deflection driver 778 is preferably a circle bridgecircuit in which power amplifiers 1446 and 1448 differentially driveboth ends of the x-deflection coil. The output voltages of the poweramplifiers 1446 and 1448 are coupled, through current sense resistors1450 and current sensor 1447, to the x-deflection coil via output lines1458 and 1460. Resistors 1450 sense the current in the x-deflection coiland preferably feeds the current information back to regulate thecontrol amplifier 1454. The current in the x-deflection coil is alsomonitored by a current sensor 1447, which transmits the measuredcurrent, via output line 1449, to the current sense monitor 788.Temperature sensor 1445, which measures the temperature at thex-deflection driver 778, employs a temperature switch 1462 to disablethe x-deflection driver 778 if a temperature fault condition occurs. They-deflection driver 782 preferably includes a similar circuit to drivethe current in the y-deflection coil.

X-step driver 780, which preferably comprises x-step ramp control switch1462, x-step voltage control circuit 1464, and decay control circuit1468 (FIGS. 57A-B), is preferably employed to generate a sawtoothcurrent wave form in the x-step coil. The x-step driver 780 is connectedacross the x-step coil via output leads 1472 and 1474. Referring to FIG.57A, x-step amplitude control signals XSTEPAMP from the beam controllerinterface 794 are preferably applied to x-step voltage control circuit1464 to control the voltage level of a VICOR multi-output switchingpower supply (not shown), which supplies an input voltage to the x-stepdriver 780 via input line 1470.

Ramp switch control signals XSTEP\ are preferably applied from thecontrol PAL 1410, via input line 1471, to control the operation of thex-step ramp control switch 1462. When the x-step ramp control switch1462 is switched on, voltage from the VICOR multi-output power supply isapplied to the x-step coil, allowing the current in the x-step coil toramp up for a specified time period, preferably 1 to 200 nsec. Theamplitude of the current pattern is determined by the voltage level ofthe VICOR multi-output power supply, which is preferably set by thex-step voltage control circuit 1464.

When the x-step ramp control switch 1462 is switched off, decay controlcircuit 1468 applies a voltage to the x-step coil to control and shapethe slope of the current decay in the x-step coil. X-step slope controlsignals XSTEPSLP are preferably applied to the decay control circuit1468 via input line 1432. An isolation amplifier 1474 is preferablyemployed to optically couple the x-step slope control signals XSTEPSLPto the decay control circuit 1468, to avoid potential problems relatingto high voltages applied to the circuit by the VICOR power supply. Theoutput of the isolation amplifier 1474 is preferably coupled to anintermediate x-step amplifier 1478. Intermediate x-step amplifier 1478preferably converts the differential output from isolation amplifier1474 into a single ended signal, which is coupled to the inverting inputof a control amplifier 1476. Control amplifier 1476 manages the voltageacross transistor 1472, which functions as a variable load, such thatthe voltage applied to the x-step coil during the current decay periodproduces an optimal current decay rate in the x-step coil. If aparticular x-ray imaging application requires the use of a y-step coil,then a y-step driver similar to the x-step driver of FIGS. 67A-B ispreferably employed.

Catheter Including X-Ray Sensitive Optical-Sensor Location Device

Referring to FIGS. 58-60, diagrams are presented of a catheter 510 thatis a presently preferred embodiment of this aspect of the invention. Asshown in FIG. 58, catheter 510 may be formed from a variety ofmaterials, such as plastic and steel. Catheter 510 comprises anelongated body 512 having a distal end 514 adapted to be inserted into abody cavity, vessel, tract, or the like and a proximal end 516 whichremains outside of the patient or object into which the catheter 510 isinserted.

FIG. 59 depicts the cross-sectional view of catheter 510. In FIG. 59,catheter 510 is shown to include a plurality of lumens 518, 520, and522. While the catheter 510 is shown including three lumens 518, 520,and 522, this embodiment is merely illustrative and the number of lumensprovided in any actual catheter fabricated according to the teachings ofthe present invention and the uses therefore will depend solely on theend use to which catheter 510 will be put. Such a catheter may besteerable to be positioned by manipulation as is known in the art.

Referring again to FIG. 58, an optical fiber 524 is disposed in lumen518 of catheter 510. For small diameter catheters, such as ones used incardiac applications, optical fiber 524 preferably has a diameter ofapproximately 0.010-0.020 inches. Those of ordinary skill in the artwill of course recognize that optical fiber 524 may have other diameterswithout departing from the scope of this aspect of the presentinvention.

An x-ray marker 526 (also referred to herein as an x-ray mini sensor),is optically coupled to a distal end 528 of the optical fiber 524.Presently preferred is an x-ray marker 526, shaped into a cube withsides of length 0.030 in. (0.075 cm), that is glued with optical cementto a 0.018 in. diameter optical fiber 524. X-ray marker 526 preferablycomprises an x-ray sensitive scintillating phosphor or crystal material,such as for example, terbium-doped gadolinium oxysulfate (Gd₂ O₂ S₂:Tb), available from USR Optronix, Inc. of Hackettstown, N.J. Thepresently preferred material for the x-ray marker 526 is a scintillatorcrystal material such as: (1) YSO (cerium doped yttriumoxy-orthosilicate, available from Airtron (Litton) of Charlotte, N.C.);(2) LSO (cerium doped lutetium oxy-orthosilicate, available fromSchlumberger, Inc.); and (3) BGO (bismuth germanate, available fromRexon Components, Inc. of Beachwood, Ohio). YSO and LSO are advantageousin that they have short decay times and therefore may be used to measurex-ray intensity by pulse counting the incident x-ray photons. Those ofordinary skill in the art will recognize that other materials or deviceswhich are responsive in the x-ray spectrum may also be used in thepresent invention and are to be considered as x-ray marker materialswithin the meaning of that phrase as used herein.

FIG. 60 diagrams the distal end of an embodiment of catheter 510according to this invention. In this embodiment, terbium dopedgadolinium oxysulfate phosphor material in a finely ground powder formis preferably both physically mounted and optically coupled to thedistal 528 end of optical fiber 524 using an epoxy 530, such asQUICKSTIK, available from GC Electronics of Rockford, Ill. The epoxy 530is preferably used to adhere the x-ray marker 526 to the optical fiber524 by placing a small amount of the epoxy 530 at the distal end ofoptical fiber 524 and then dipping the distal end 528 of optical fiber524 into a finely divided phosphor powder. Those of ordinary skill inthe art will recognize that the epoxy 530 is preferably compatible withthe phosphor material and is preferably optically clear at the phosphoremission wavelength. Similar techniques may be employed to affix ascintillating crystal to the distal end 528 of optical fiber 524. Inaddition, other methods for optically coupling a marker material to anoptical fiber, such as employment of a lens, are equivalent and fallwithin the scope of the present invention.

A photodetector 532 is optically coupled to the proximal end 534 ofoptical fiber 524. Photodetector 532 may be a photodiode, aphotomultiplier tube, a phototransistor, a charge-coupled device, orother device which transduces the light signal generated at x-ray marker526 due to exposure to an x-ray flux into an electrical signal when thelight signal is transmitted down optical fiber 524 to photodetector 532.The magnitude of the electric signal generated by photodetector 532 isrelated to the magnitude of the intensity of the x-ray flux sensed byx-ray mini sensor 526. In one preferred embodiment of this invention,the proximal end 534 of optical fiber 524 is coupled to thelight-sensitive window of an RCA XT2020 photomultiplier tube using ahigh viscosity clear silicone oil, such as available from Dow CorningCorporation of Midland, Mich.

The electrical signal from photodetector 532 is converted to a digitalsignal by conventional means such as an A/D converter and then brokendown into a series of values representing the instantaneous x-ray fluxintensity at x-ray mini sensor 526 corresponding to the variousapertures 140 of collimation grid 90--thus as the apertures 140 ofcollimation grid 90 sequentially emit x-ray pencil beams 50, a valuecorresponding to the measured x-ray flux intensity at sensor 526 isstored for each aperture in a memory array for use as discussed below.

As will be discussed in further detail below, an electrical signal willbe generated by photodetector 532 when the catheter 510 is exposed tox-ray radiation. Catheter 510 of the present invention is thereforehighly useful when employed in applications with reverse-geometryscanning-beam x-ray systems, enabling location of the catheter'sposition with a high degree of accuracy.

Although the catheter 510 of FIGS. 58-60 is shown employing a singleoptically-coupled x-ray marker 526, catheters having a plurality of suchoptically coupled x-ray markers may be fabricated according to theprinciples of the present invention. The number of such markers employedin any actual embodiment of the catheter of the present invention willdepend solely on the application for which the particular catheter isdesigned. Specific illustrative examples are disclosed herein.

This aspect of the present invention may be employed in a wide varietyof both steerable and non-steerable catheters for various applications.Without limitation, such catheters and applications will includemulti-electrode catheters employed in electrophysiology, and cathetersfor ablation of cardiac arrhythmia, coronary atherectomy catheters, etc.This aspect of the present invention will also be usefully employed, forexample, in angioplasty balloon catheters, and in laparoscopy equipment.Such catheters now employ bands of radio-opaque material such as metal,so that they may be located by x-ray techniques during performance ofangioplasty procedures. Those of ordinary skill in the art willdoubtless be able to envision other applications for the presentinvention.

Referring now to FIGS. 61a-61c, certain exemplary catheters employingthe present invention are described. Referring first to FIG. 61a, amulti-electrode electrophysiology catheter 540 for use in measuringintracardiac action potentials incorporates an embodiment of the presentinvention and is shown diagrammatically. An example of such a commercialelectrophysiology catheter of this type is Steerocath Model No. 2037available from EP Technologies, Inc. of Mountain View, Calif. Forconvenience, features of the catheter 540 which are the same as featuresof catheter 510 of FIGS. 58-60 are identified by the same referencenumerals. Thus, catheter 540 is shown to include a plurality of lumens518, 520, and 522. An optical fiber 524 is disposed in lumen 518 ofcatheter 540. An x-ray marker material 526 is affixed to a distal end528 of the optical fiber 524.

A plurality of conductive electrodes 542, 544, 546, and 548 are disposedin spaced-apart relationship near the distal end 514 of catheter 540.Electrodes 542, 544, 546, and 548 are individually electrically coupledto conductors 550, 552, 554, and 556, respectively. Conductors 550, 552,554, and 556 are disposed in lumen 520 and emerge from its orifice atthe proximal end 516 of catheter 540 and may be suitably connected to anelectrophysiology recording and analysis system as is known in the art.As shown in FIG. 61a, a photodetector 532 is coupled to the proximal end534 of optical fiber 524.

Referring now to FIG. 61b, a diagram of a balloon angioplasty catheter560 employing an embodiment of the present invention is presented. Forconvenience, features of the angioplasty catheter 560 which are the sameas features of catheter 510 of FIGS. 58-60 are identified by the samereference numerals. Thus, angioplasty catheter 560 comprises anelongated body 512 having a distal end 514 adapted to be inserted into abodily cavity, vessel, tract, or the like and a proximal end 516available to a person performing a medical procedure.

Angioplasty catheter 560 includes an inflatable balloon 562 disposednear its distal end 514 which is used as is known in the art. Balloon562 may be inflated from a lumen 564 with which it communicates.Angioplasty catheter 560 preferably includes a central guidewire lumen566 through which a guidewire is threaded as is known in the art.

Angioplasty catheter 560 according to the present invention hasadvantages over prior art angioplasty catheters because it incorporatesa plurality of x-ray markers as taught by the present invention. Theembodiment illustrated in FIG. 61b shows three such x-ray markers 526a,526b, and 526c, but those of ordinary skill in the art will recognizethat an arbitrary number of markers could be employed if needed.

In the particular example of FIG. 61b, x-ray marker 526a is disposednear the distal side of the balloon and x-ray marker 526b is disposednear the proximal side of the balloon. In addition, x-ray marker 526c isdisposed further down the body 512 of the catheter towards its proximalend 516. X-ray markers 526a, 526b, and 526c are optically coupled tofirst ends of optical fibers 524a, 524b, and 524c, respectively. Opticalfibers 524a, 524b, and 524c are shown disposed in individual lumens518a, 518b, and 518c, although those of ordinary skill in the art willrecognize that they might all be disposed in a single lumen if designconsiderations warranted such placement.

Optical fibers 524a, 524b, and 524c are connected at their second endsto photodetectors 532a, 532b, and 532c. The present invention representsa significant advantage over prior art catheters in which x-ray opaquebands are physically attached to selected positions of the catheter. Thecatheter of the present invention provides the advantage that thepositions of the individual x-ray markers (and thus the position ofballoon 562) may be uniquely and accurately determined due to theirindividual coupling to photodetectors 532a, 532b, and 532c. The presentinvention is easily interfaced to the widely used digital computer videoequipment employed with modern x-ray systems so that, when used inconjunction with a scanning x-ray system, the positions of the x-raymarkers may be indicated, for example, by blinking cursors, etc.,avoiding the ambiguity which is characteristic of interpretation ofx-ray displays to locate prior-art catheters.

Referring now to FIG. 61c, a diagram of an atherectomy catheteremploying an embodiment of the present invention is presented. Forconvenience, features of the catheter 580 which are the same as featuresof catheter 510 of FIGS. 58-60 are identified by the same referencenumerals. Thus, atherectomy catheter 580 preferably comprises anelongated body 512 having a distal end 514 adapted to be inserted into abody cavity, vessel, tract, or the like and a proximal end 516 availableto a person performing a medical procedure.

A central guidewire lumen 582 runs through the body of atherectomycatheter 580 from distal end 514 to proximal end 516 as is known in theart. A tapered coil spring 584 is typically internally disposed at thedistal end 514 of atherectomy catheter 580. A rotating cutting section586 comprises a cylindrical member axially disposed in the catheter bodybehind the tapered distal end 514. A cutting blade 588 is slidablymounted in rotating cutting section 586. Rotating cutting section 586 iscoupled to a shaft 590 which runs through a lumen 582 and may be coupledto a motor for rotation. These elements and their relationship to oneanother are well known.

Atherectomy catheter 580 has advantages over known atherectomy cathetersbecause it incorporates a plurality of x-ray markers as taught by thepresent invention. The embodiment illustrated in FIG. 61c shows threesuch x-ray markers 526a, 526b, and 526c, but those of ordinary skill inthe art will recognize that an arbitrary number of markers could beemployed if needed.

In the particular example of the atherectomy catheter of FIG. 61c, x-raymarker 526a is disposed near the distal side of the rotating cuttingsection 586 and x-ray marker 526b is disposed near the proximal side ofthe rotating cutting section 586. In addition, x-ray marker 526c isdisposed further down the body 512 of the catheter towards its proximalend 516. X-ray markers 526a, 526b, and 526c are optically coupled tofirst ends of optical fibers 524a, 524b, and 524c, respectively. Opticalfibers 524a, 524b, and 524c are shown disposed in individual lumens518a, 518b, and 518c, although they might all be disposed in a singlelumen if design considerations warranted such placement.

Optical fibers 524a, 524b, and 524c are connected at their second endsto photodetectors 532a, 532b, and 532c. The present invention representsa significant advantage over prior art catheters in which x-ray opaquebands are physically attached to selected positions of the catheter. Thecatheter of the present invention provides the advantage that thepositions of the individual x-ray markers (and thus of rotating cuttingsection 586) may be uniquely and accurately determined due to theirindividual coupling to photodetectors 532a, 532b, and 532c. As in theembodiment of FIG. 61b, the positions of the x-ray markers inatherectomy catheter 580 of FIG. 61c may be indicated, for example, byblinking cursors, etc., avoiding the ambiguity which is characteristicof interpretation of x-ray displays to locate prior-art catheters.

In the case of the atherectomy catheter, it is important to know whichway the opening for the cutting blade is facing. This can be readilydetermined according to the present invention by partially shielding thex-ray marker material 526, for example, on three sides, leaving onlyapproximately 90° of rotation open to radiation. If more than one markeris shielded and the openings are oriented in different directions, anaccurate assessment of catheter rotational orientation may be made.

This aspect of the present invention is illustrated in FIG. 61d, whichis a view down the catheter of a plurality of x-ray markers according tothe present invention. Although four x-ray markers are depicted in FIG.61d, other numbers may be employed to affect the resolution with whichthe rotational position of the catheter may be determined.

Referring now to FIG. 61d, x-ray markers 526a, 526b, 526c, and 526d aredisposed in a shielding member 590, which is illustrated in crosssection as generally cross shaped. Shielding member 590 is fabricatedfrom a material which will attenuate x-ray radiation. In the typicalmedical x-ray environment, shield member 590 may comprise, for example,tungsten having a thickness of about 0.010 in. It will be possible toutilize other materials and thicknesses because of the ability tocompare the outputs of different photodetectors.

Each of x-ray markers 526a, 526b, 526c, and 526d is separately connectedto an optical fiber and a photodetector (not shown). In theconfiguration shown the rotational position of the catheter can bedetermined by comparing the outputs of the photodetectors.

Those of ordinary skill in the art will recognize that the FIG. 61d doesnot indicate whether x-ray markers 526a, 526 b, 526c, and 526d aredisposed at the same point along the length of the catheter in whichthey are placed. This does not matter and will be a matter of designchoice, subject of course to the inaccuracy which would result fromtwisting of the catheter if they are disposed at different positionsalong the length of the catheter.

The particular applications and examples disclosed herein are merelyillustrative and those of ordinary skill in the art will be readily ableto envision other applications for which the present invention issuited. It is intended that all such applications fall within the scopeof the present invention.

Determination of the X and Y Coordinate Location of a ManeuverablePositioner

Referring to FIG. 62, this diagram illustrates one technique to locatethe precise x and y coordinates of a maneuverable positionerincorporating an x-ray marker or x-ray mini sensor. The x and ycoordinates specify the location where the x-ray marker 526 of amaneuverable positioner is located, on a plane 280 which is parallel tothe output face 260 of collimation grid 90. In FIG. 62, a preferredreverse-geometry scanning-beam x-ray imaging system 600 is shownemployed with catheters 510a and 510b disposed in an object 602. Object602 could be, for example, a patient into which catheters 510a and 510bhave been inserted. Photodetectors 532a and 532b are coupled to theproximal ends of the optical fibers inside of catheters 510a and 510brespectively. Two catheters 510a and 510b, and their associatedphotodetectors 532a and 532b are shown for purposes of illustration, butthose of ordinary skill in the art will recognize that any number ofcatheters and photodetectors may be employed according to the principlesof the present invention.

Collimation grid 90 is preferably positioned between the x-ray tube 604,and an x-ray detector 610 to ensure that x-ray beams from selectedpositions of the scanned x-ray tube 604 are directed to the x-raydetector 610. As the x-ray beams travel their various paths from thecollimation grid 90 to the x-ray detector 610, some of the emissions areintercepted by and irradiate the x-ray sensors 526a and 526b disposed onthe distal ends of optical fibers 524a and 524b in catheters 510a and510b, respectively. The x-ray markers respond by emitting light, whichis directed down the optical fibers 524a and 524b to photodetectors 532aand 532b, which convert the light to electrical signals, which arefurther conditioned and amplified before being transmitted to a controlunit 620.

Control unit 620 is supplied with the x and y addresses from the scancontroller 606 and with the output signals from photodetectors 532a and532b. Control unit 620 correlates the output signals from photodetectors532a and 532b with the scan address information from scan controller606. This information is processed to determine the positions of thex-ray sensor in catheters 510a and 510b, and hence is an accurateindicator of the positions of the ends of the catheters themselves.

FIG. 63 diagrams three preferred x-ray detection methods for catheterlocation. For the binary detection method, the control unit 620 employsan integrator and a comparator to process the amplified output signalsfrom the photodetectors 532a and 532b, to determine if the x-ray sensors526a and 526b were illuminated by x-rays for each x and y scan address.

The energy measurement method employs a integrator and ananalog-to-digital converter to measure the x-ray energy levels receivedat the x-ray sensors 526a and 526b for each x and y scan address. Inthis method, the x-ray intensity values at x-ray sensors 526a and 526bof catheters 510a and 510b respectively, are measured for each x-raybeam that is emitted at the target. Since each aperture of collimationgrid 90 may be scanned once per x-ray tube scan cycle (this need not bethe case and any predetermined group of apertures could be used as longas they reasonably cover the field), the measured x-ray intensity valuesfrom the x-ray detector 610 may be stored in a memory array associatedwith a computer processor. FIG. 64 shows the sample data which may beaccumulated within such a memory array during one scan cycle of thex-ray tube for a single x-ray marker. The vertical axis 626 of FIG. 64represents x-ray flux intensity as measured by x-ray marker 526 and thehorizontal axes 628 and 630 represent the X, Y coordinates of theposition respectively, of collimator aperture 140 which was beingilluminated when the intensity data was captured.

As is shown, there may be spurious intensity readings scatteredthroughout the array. These are marked with she numeral 632. Thespurious readings 632 are most likely the result of scattered x-raysstriking the x-ray mini sensor 526 and not of primary x-rays emanatingdirectly from collimation grid 90. The spurious readings 632 arepreferably filtered out by well-known techniques leaving the main bodyof the data which is depicted here as a large tall structure 634. Thisstructure 634 represents mostly x-rays directly from apertures ofcollimation grid 90 which were measured by x-ray marker 526. The shapeand width of this structure will vary depending upon the distance Z ofx-ray mini sensor 526 from target 50. It is also possible to determine Zfrom this information alone, as detailed below. The X and Y coordinatesof the position of structure 634 will vary within the data arraydepending upon the X and Y coordinates of the position of x-ray marker526 in space. The actual X and Y coordinates of the position of x-raymarker 526 can be determined by analyzing this data in a conventionalmanner to identify the collimation grid aperture which lies at itsgeometric center or centroid. Then, given the distance Z determinedbelow, the x-ray marker 526 lies on a plane a distance Z from target 50and on a line from target 50 to x-ray detector 610 which passes throughthe identified aperture. This defines a precise position in space whichcan be easily calculated.

The final method is the photon counting method, which employs acomparator and counter to determine the number of photons detected atthe x-ray sensors 526a and 526b for each x and y scan address. For thismethod, the x-ray sensors 526a and 526b are preferably formed of ascintillator material with a fast response time, e.g., YSO and LSO, toprovide a signal train which can be used to count x-ray photonsimpinging upon the x-ray sensor. The amplified output signals from thephotodetectors 532a and 532b are fed to a comparator, which outputs asignal pulse when the input signal rises above a threshold correspondingto individual x-rays striking the x-ray mini sensor 526. Data from thephoton counting method is preferably processed in the same way as forthe energy measurement method above.

Determination of the Z Coordinate of a Maneuverable Positioner

The presently preferred method to determine the Z-axis coordinate (orheight) of a maneuverable positioner which incorporates an x-ray markeris to count the number of x-ray pencil beams that strike the x-raymarker per image frame. The Z coordinate or height value of a givenposition is normally a distance from target 50 to a plane 280 parallelto output face -260 of collimator grid 90 in which sensor 526a islocated, although it can obviously be linearly translated to be adistance from any reference point to x-ray marker 526a. To illustratethe presently preferred method, references will be made to the catheter510a and scanning-beam x-ray system 600 as diagrammed in FIG. 62.

As target 50 is scanned by an electron beam 40, x-rays are emitted invirtually all directions at the surface of target 50. Referring to FIG.65, collimation grid 90 is preferably fabricated as discussed above of arelatively x-ray opaque material and has a plurality of apertures 140disposed therein which provide paths for the x-rays from collimationgrid input face 66 to collimation grid output face 260 along which someof these x-rays may travel. X-rays which emanate through apertures 140of collimation grid 90 preferably form into x-ray pencil beams 100.

As shown in FIG. 65, these individual x-ray pencil beams 100 divergeslightly into a cone shape 624 as they proceed toward x-ray detector610. It should be remembered that only one x-ray pencil beam 100 will begenerated at any one time. FIG. 65, for purposes of illustration, onlydepicts 8 x-ray pencil beams 100. At plane A1-A2, which cuts through thex-ray pencil beams 100 parallel to collimation grid output face 260, thex-ray beams have not diverged enough to create beam overlaps. Thus asmall x-ray marker 526a located within plane A1-A2 will "see" only asingle beam at a time. FIG. 66A depicts the x-ray pencil beam 100cross-sectional pattern at plane A1-A2 for a 7-by-7 array of beams.Moving away from collimation grid 90 is plane B1-B2 which is parallel toplane A1-A2. FIG. 66B similarly shows the x-ray pencil beam 100cross-sectional pattern at plane B1-32. In plane B1-B2, the x-ray pencilbeams 100 have diverged enough to create some overlap among the x-raypencil beams 100. Thus a small x-ray marker 526a located at plane B1-B2may "see" more than a single beam. The diagrams of FIGS. 66C and 66D ofplanes C1-C2 and D1-D2 respectively show what happens as a small x-raysensor 526 approaches x-ray detector 610--it "sees" more and moreoverlapping x-ray pencil beams 100 the closer it gets to the x-raydetector 610. At the surface of x-ray detector 610 it will see all ofthe x-ray- pencil beams 100 emanating from collimation grid output face260.

As discussed above, the pencil beams of x-rays 100 are not all generatedsimultaneously. Because the x-ray tube is a scanning x-ray tube--muchlike the cathode ray tube used in television sets and computermonitors--at most one aperture is "illuminated" at any given instant byelectron beam 40 and, as a result, generates x-rays which emanate fromthe aperture 140 corresponding to that location on target 50 at thatinstant. Thus, to an x-ray marker 526 within the cone 624 of x-raysformed between collimation grid output face 260 and x-ray detector 610,a series of x-ray pulses will be seen during each cycle of the x-raytube (a "cycle" of the x-ray tube corresponds to a complete scan andtypically occurs, in a preferred embodiment of the present invention,15-30 times per second depending on how the system is adjusted by theuser). This series of pulses corresponds to the x-ray pencil beams lowhich are seen by the x-ray sensor 526. Accordingly, an x-ray marker 526located close to the output face 260 of collimation grid 90 will measurea relatively low number of pulses per x-ray tube cycle while the samex-ray marker, if located close to x-ray detector 610 will measure a muchhigher number of pulses per x-ray cycle. It is this property thatpermits the measurement of the Z coordinate.

A method for determining the Z coordinate based upon the number ofpulses per x-ray tube cycle is derived mathematically as follows:

It is assumed- that the beam diameter d_(b) is 0 at target 50 and d_(c)at the collimation grid output face 260. For any distance Z from target50: ##EQU9## The spacing between the x-ray pencil beams 100 λ_(b) varieslinearly with λ_(s) at target 50 to 0 at the x-ray detector 610. For anydistance Z from target 50: ##EQU10## A square unit of area within thecone-shaped x-ray field 624 generated by the x-ray system has a numberof beam centers passing through it of: ##EQU11## The total area of thesebeams is: ##EQU12## if there is no overlap. Since the beams fit into asquare of unit area, the overlap is: ##EQU13## Assuming that λ_(s)=0.0203 in (0.052 cm), d_(c) =0.015 in (0.0380 cm), Z_(c) =1.0 in (2.54cm) and Z_(d) =37.2 in (94.5 cm), then the overlap at a given distance Zis set forth in the following TABLE III:

                  TABLE III    ______________________________________    DISTANCE (in)    (cm)   OVERLAP    ______________________________________    1                2.54   0    2                5.08   2    3                7.62   5    4                10.16  9    5                12.70  15    6                15.24  23    7                17.78  33    8                20.32  46    9                22.86  62    10               25.40  83    11               27.94  108    12               30.48  139    13               33.02  17    14               35.56  223    15               38.10  279    16               40.64  348    17               43.18  433    18               45.72  537    19               48.26  666    20               50.80  827    21               53.34  1027    22               55.88  1281    23               42     1604    24               60.96  2021    25               63.50  2567    26               66.04  3295    27               8.58   4284    28               71.12  5663    29               73.66  7647    30               76.20  10614    ______________________________________

Thus, for example an x-ray sensor 526a in catheter 510a will measure 62pulses per x-ray tube scan cycle at a distance of 9 in (22.86 cm) fromtarget 50.

Another preferred method for determining the information in Table III isto measure the actual number of pulses detected by the marker 526a atpredetermined locations. One can construct a look-up table which can beconsulted to display the Z coordinate as a function of the number ofpulses detected by the marker 526a. Where possible, this method can alsobe used within a living patient or other object where it is possible todetermine a first absolute Z coordinate and a second absolute Zcoordinate within the x-ray cone 624 by another means, such as by directmeasurement, then Z coordinates within the object or patient can bedetermined by interpolation.

Another technique to measure the Z-axis or distance of the marker 526ain a catheter 510 is based upon the measurement of the x-ray intensityat the marker 526a. This technique is related to the intensitymeasurement technique to determine the x and y coordinates discussedabove. As mentioned above, when the collimator grid apertures 140 areilluminated by the x-ray beam, the x-ray mini sensor 638 from catheter510 will receive the greatest x-ray intensity from the aperture whoseaxis it is closest to. This not only gives X and Y location, but alsogives Z location in the following way. Referring to FIG. 64, the girthof structure 634 is directly related to the Z location of the marker526. If the marker 526 is far from the source, the girth will be large;if close, it will be small. The preferred place to measure the girth isakin to the full width at half maximum, i.e., measure the girth at onehalf the height of the structure 634. It can also be measured at otherpoints and will give like results. Most reliable is believed to be therange from approximately 1/4 the height to 3/4 the height.

Referring now to FIG. 66, this diagram depicts an alternate method todetermine the location of a catheter in three dimensions. Two x-raytubes 604a and 604b, employing independent scan controllers 106a and106b and collimation grids 90a and 90b, may be used in combination withcatheters 510a, 510b, and 510c, photodetectors 532a, 532b, and 532c, andcontrol unit 620 to accurately locate the positions of the catheters510a, 510b, and 510c in two dimensions with respect to each of the twosources. Triangulation can be used to combine the two 2-dimensionalmeasurements into a single 3-dimensional position.

An alternate method to determine the Z coordinate of a catheter is basedupon photon counting, i.e., by using an x-ray mini sensor 526 having afast response time, e.g., YSO and LSO, to provide a signal train whichcan be used to count x-ray photons impinging upon the x-ray sensor.Since photon counting is just another x-ray intensity measurementmethod, it can be employed with any of the foregoing methods todetermine the Z coordinate of a catheter.

Display of X, Y and Z Coordinates for Maneuverable Positioner

Given the X, Y and Z coordinates of the position of x-ray sensor 526 ascalculated above, a number of useful things can be done with that data.One preferred use is to draw a cursor on the monitor screen which isdisplaying the x-ray image of the object or patient. This can enhanceand emphasize the display of the location of x-ray sensor 526 (oftenlocated at the tip of a maneuverable positioner being positioned withinthe body or object). Similarly, "waypoints" which are defined by the X,Y and Z coordinates of positions where the sensor has been or where itis going may be displayed as computer-generated icons on the monitorscreen to guide the user in positioning the x-ray sensor and itsassociated maneuverable positioner. Waypoints may be captured and storedfor display or study and they can be captured about the surface of anobject of interest within a body or object such as an aneurysm, a stent,a tumor, features on a heart, and the like. The locus of these waypointsmay then be displayed on a computer screen to give a 2 or a 3dimensional image of the object, to study any changes in its location,condition or shape, and the like.

Apart from visual display of the data in terms of a visual image on amonitor screen, the actual X, Y and Z data numbers corresponding to alocation of interest within the patient or object may be displayed orprinted out or otherwise made available to a user so that the user mayreturn the maneuverable positioner to an exact same location. Forexample, certain procedures on the heart require that electricalmeasurements be made of the heart. The electrical values so measured maybe stored in concert with the exact locations at which the respectivemeasurements were made. Then, once it is decided how to proceed giventhe measured data, corrective action can be taken at specified siteswhich can be precisely located by means of the recorded X, Y and Zinformation corresponding to the site(s) of interest.

While embodiments, applications and advantages of the invention havebeen shown and described with sufficient clarity to enable one skilledin the art to make and use the invention, it would be equally apparentto those skilled in the art that many more embodiments, applications andadvantages are possible without deviating from the inventive conceptsdisclosed and described herein. The invention therefore should only berestricted in accordance with the spirit of the claims appended heretoand is not to be restricted by the preferred embodiments, specificationor drawings.

What is claimed is:
 1. An x-ray collimation assembly comprising:a cooledtarget comprising an x-ray generating layer and a cooling fluid; and,said cooling fluid adjacent to said x-ray generating layer, said cooledtarget comprising a temperature sensor past which said cooling fluidflows to cool a collimator; said collimator arranged such that saidcooling fluid flows between said x-ray generating layer and saidcollimator, said collimator comprising a plurality of x-ray transmissivepassages extending through said collimator, each of said x-raytransmissive passages comprising an axis, and said x-ray transmissivepassages arranged to collimate x-rays generated by said target; saidcollimator comprising a substantially planar configuration, the axis ofeach of said x-ray transmissive passages forming an angle with the planeof said collimator and substantially converging at a predetermined area.2. The collimation assembly of claim 1 wherein said collimator comprisesa first endplate adjacent a first surface of said collimator andcomprised of an x-ray transmissive material.
 3. The collimation assemblyof claim 1 wherein said collimator comprises a first endplate adjacent afirst surface of said collimator and a second endplate adjacent a secondsurface of said collimator wherein each of said endplates are comprisedof an x-ray transmissive material.
 4. The collimation assembly of claim1 wherein said x-ray transmissive passages are formed by chemicaletching.
 5. The collimation assembly of claim 1 wherein said x-raytransmissive passages are formed by one of the machining processes fromthe group consisting of electron beam machining, drilling,mini-machining and laser drilling.
 6. The collimation assembly of claim1 wherein said x-ray transmissive passages are cylindrically shaped. 7.The collimation assembly of claim 1 wherein said collimator is comprisedof at least one of the materials selected from the group consisting of:brass, tungsten, lead, molybdenum.
 8. The collimation assembly of claim1 wherein said cooled target further comprises a target supportcomprised of an x-ray transmissive material.
 9. A collimation assemblyof claim 8 wherein said target support is comprised of beryllium andsaid x-ray generating layer is comprised of tantalum or tungsten. 10.The collimation assembly of claim 8 wherein an intermediate layer of aresilient material is disposed between said target support and saidx-ray generating layer.
 11. The collimation assembly of claim 10 whereinsaid resilient material is niobium.
 12. The collimation assembly ofclaim 1 wherein said collimator comprises a plurality of sheets of x-rayabsorbing material.
 13. The collimation assembly of claim 1 wherein saidx-ray transmissive passages are holes.
 14. An x-ray target apparatuscomprising a first surface;a target assembly forming said first surface;a cooling medium; and, a grid, said grid comprising a plurality of x-rayabsorbent sheets, each comprising a plurality of x-ray transmissiveapertures said plurality of x-ray absorbent sheets arranged one atop theother to form a substantially planar stack; said x-ray transmissiveapertures of sheets aligned with the x-ray transmissive apertures ofother sheets to form an x-ray transmissive passage through said stack;the axis of each of said x-ray transmissive passages substantiallyconverging, said stack permanently secured at a periphery of saidsheets; said target assembly comprising a target support and an x-rayemitting layer, and said cooling medium arranged to cool said grid andsaid target such that said temperature of said first surface isapproximately 40 degrees C.
 15. The x-ray target apparatus of claim 14further comprising alignment structures precisely located on at leastsome of said sheets.
 16. The x-ray target apparatus of claim 15 whereinsaid x-ray target apparatus further comprises a second surface andwherein said coolant is arranged to cool said second surface through acoolant chamber.
 17. The x-ray target apparatus of claim 14 wherein saidx-ray transmissive areas are formed by chemical etching of the sheets.18. The x-ray target apparatus of claim 14 wherein the x-raytransmissive apertures are cylindrically shaped.
 19. The x-ray targetapparatus of claim 14 wherein said sheets comprise at least one of thematerials selected from the group consisting of: brass, tungsten, lead,molybdenum.
 20. The x-ray target apparatus of claim 14 wherein saidtarget support material is beryllium and said x-ray emitting layer iscomprised of tungsten or tantalum.
 21. The x-ray target apparatus ofclaim 14 wherein said x-ray target apparatus further comprises a secondsurface and wherein said coolant is also arranged to cool said secondsurface to a temperature less than the melting point of said x-ray layeremitting material.
 22. An x-ray grid assembly comprisinga plurality ofx-ray absorbent sheets, each comprising a plurality of x-raytransmissive areas; said plurality of x-ray absorbent sheets arrangedone atop the other to form a substantially planar stack; said x-raytransmissive areas of each sheet aligned with the x-ray transmissiveareas of the immediately adjacent sheets to form a first set and asecond set of x-ray transmissive passages through said stack; the axisof each of said x-ray transmissive passages of said first set forming anangle ranging from 40° to 90° with the plane of said stack andsubstantially converging at a first spot, and the axis of each of saidx-ray transmissive passages of said second set forming an angle rangingfrom 40° to 90° with the plane of said stack and substantiallyconverging at a second spot; a target and a coolant chamber wherein saidcoolant chamber is disposed adjacent said stack and said target iscomprised of a target support and an x-ray emitting layer.
 23. The gridassembly of claim 22 wherein the x-ray transmissive areas of said firstset and said second set are cylindrically shaped.
 24. An x-ray gridassembly comprisinga first plurality of x-ray absorbent sheets, eachcomprising a plurality of x-ray transmissive areas arranged in a regularpattern, said first plurality of x-ray absorbent sheets comprised of amaterial having a high atomic number; a second plurality of x-rayabsorbent sheets, each comprising a plurality of x-ray transmissiveareas, said second plurality of x-ray absorbent sheets comprised of amaterial having a low atomic number; said first plurality and secondplurality of x-ray absorbent sheets arranged one atop the other to forma substantially planar stack; said x-ray transmissive areas of eachsheet aligned with the x-ray transmissive areas of the immediatelyadjacent sheets to form an x-ray transmissive passage through saidstack; the axis of each of said x-ray transmissive passages forming anangle ranging from 40° to 90° with the plane of said stack andsubstantially converging at a single spot; a target and a coolantchamber wherein said coolant chamber is disposed adjacent said stack andsaid target is comprised of a target support and an x-ray emittinglayer.
 25. The grid assembly of claim 24 wherein the x-ray transmissiveareas are cylindrically shaped.
 26. The grid assembly of claim 24wherein said first plurality of sheets comprise at least one of thematerials selected from the group consisting of: tungsten, lead,molybdenum.
 27. The grid assembly of claim 24 wherein the material forsaid second plurality of sheets is brass.
 28. The grid assembly of claim24 wherein said target support material is beryllium and said x-rayemitting layer is comprised of tantalum or tungsten.